System and process of utilizing energy for treating biological tissue

ABSTRACT

A process for heat treating biological tissue includes providing a plurality of energy emitters formed into an array. Treatment energy is generated from the plurality of emitters and applied to target tissue. The treatment energy has energy and application parameters selected so as to raise the target tissue temperature sufficiently to create a therapeutic effect while maintaining an average temperature of the target tissue over several minutes at or below a predetermined temperature so as not to destroy or permanently damage the target tissue.

RELATED APPLICATIONS

This application is a continuation-in-part of U.S. application Ser. No.15/918,487 filed Mar. 12, 2018, which is a continuation-in-part of U.S.application Ser. No. 15/629,002 filed Jun. 21, 2017, Ser. No. 15/583,096filed May 1, 2017, Ser. No. 15/460,821 filed Mar. 16, 2017, Ser. No.15/232,320 filed Aug. 9, 2016 (now U.S. Pat. No. 9,962,291), Ser. No.15/214,726 filed Jul. 20, 2016, Ser. No. 15/178,842 filed Jun. 10, 2016(now U.S. Pat. No. 9,626,445), Ser. No. 14/922,885 filed Oct. 26, 2015(now U.S. Pat. No. 9,427,602), Ser. No. 14/921,890 filed Oct. 23, 2015(now U.S. Pat. No. 9,381,116), Ser. No. 14/607,959 filed Jan. 28, 2015(now U.S. Pat. No. 9,168,174), Ser. No. 13/798,523 filed Mar. 13, 2013,and Ser. No. 13/481,124 filed May 25, 2012. This application is also acontinuation-in-part of U.S. application Ser. No. 15/460,821, filed Mar.16, 2017.

BACKGROUND OF THE INVENTION

The present invention is generally directed to systems and processes fortreating biological tissue, such as diseased biological tissue. Moreparticularly, the present invention is directed to a process for heattreating biological tissue using energy having parameters and appliedsuch so as to create a therapeutic effect to a target tissue withoutdestroying or permanently damaging the target tissue.

The inventors have discovered that there is a therapeutic effect tobiological tissue, and particularly damaged or diseased biologicaltissue, by controllably elevating the tissue temperature up to apredetermined temperature range while maintaining the averagetemperature rise of the tissue over several minutes at or below apredetermined level so as not to permanently damage the target tissue.More particularly, the inventors have discovered that electromagneticradiation, such as in the form of various wavelengths of light, can beapplied to retinal tissue in a manner that does not destroy or damagethe retinal tissue while achieving beneficial effects on eye diseases.The inventors have found that a light beam can be generated and appliedto the retinal tissue cells such that it is therapeutic, yet sublethalto retinal tissue cells and thus avoids damaging photocoagulation in theretinal tissue which provides preventative and protective treatment ofthe retinal tissue of the eye. The treatment typically entails applyinga train of laser micropulses to radiate a portion of a diseased retinafor a total duration of less than a second. Each micropulse is on theorder of tens to hundreds of microseconds long, with the microsecondsbeing separated by one to several milliseconds, which raises the tissuetemperature in a controlled manner.

It is believed that raising the tissue temperature in such a controlledmanner selectively stimulates heat shock protein activation and/orproduction and facilitation of protein repair, which serves as amechanism for therapeutically treating the tissue. It is believed thatthis micropulse train thermally activates heat shock proteins (HSPs) inthe targeted tissue. In the case of retinal tissue, the processthermally activates HSPs in the retinal pigment epithelium (RPE) layerimmediately behind the retinal layer containing the visually sensitiverods and cones, and that these activated HSPs then reset the diseasedretina to its healthy condition by removing and repairing damagedproteins. This then results in improved RPE function, improves retinalfunction and autoregulation, restorative acute inflammation, reducedchronic inflammation, and systematic immunodulation. Theselaser-triggered effects then slow, stop or reverse retinal disease,improve visual function and reduce the risk of visual loss. It isbelieved that raising tissue temperature in such a controlled manner toselectively stimulate heat shock protein activation has benefits inother tissues as well.

HSPs are a family of proteins that are produced by cells in response toexposure to stressful conditions. Production of high levels of heatshock proteins can be triggered by exposure to different kinds ofenvironmental stress conditions, such as infection, inflammation,exercise, exposure of the cell to toxins, oxidants, heavy metals,starvation, hypoxia, water deprivation and tissue trauma.

It is known that heat shock proteins play a role in responding to alarge number of abnormal conditions in body tissues, including viralinfection, inflammation, malignant transformations, exposure tooxidizing agents, cytotoxins, and anoxia. Several heat shock proteinsfunction as intra-cellular chaperones for other proteins and members ofthe HSP family are expressed or activated at low to moderate levelsbecause of their essential role in protein maintenance and simplymonitoring the cell's proteins even under non-stressful conditions.These activities are part of a cell's own repair system, called thecellular stress response or the heat-shock response.

Heat shock proteins are found in nearly every cell and tissue-type ofmulticellular organisms as well as in explanted tissues and in culturedcells. The HSPs typically comprise 3%-10% of a cell's proteins, althoughwhen under stress the percentage can rise to 15%. The density ofproteins of a mammalian cells has been found to be in the range of(2-4)×10¹⁸CM⁻³. Thus, the aforementioned percentages mean that thedensity of HSPs is normally (1-4)×10¹⁷CM⁻³, while under stress thedensity can rise to (3-6)×10¹⁷CM⁻³.

Heat shock proteins are typically named according to their molecularweight, and act in different ways. An especially ubiquitous heat shockprotein is Hsp70, a protein with a molecular weight of 70 killodaltons.It plays a particularly significant role in protecting proteins that arejust being formed and in rescuing damaged proteins. It contains a groovewith an affinity for neutral, hydrophobic amino acid residues that caninteract with peptides up to 7 residues in length. Hsp70 haspeptide-binding and ATPase domains that stabilize protein structures inunfolded and assembly-competent states. The HSPs play a role inpreventing aggregation of misfolded proteins, many of which have exposedhydrophobic portions, and a facilitating the refolding of proteins intotheir proper conformations. Hsp70 accomplishes this by first binding tothe misfolded or fragmented protein, a binding that is madeenergetically possible by a site that binds ATP and hydrolyzes it intoADP.

Hsp70 heat shock proteins are a member of extracellular and membranebound heat-shock proteins which are involved in binding antigens andpresenting them to the immune system. Hsp70 has been found to inhibitthe activity of influenza A virus ribonucleoprotein and to block thereplication of the virus. Heat shock proteins derived from tumors elicitspecific protective immunity. Experimental and clinical observationshave shown that heat shock proteins are involved in the regulation ofautoimmune arthritis, type 1 diabetes, mellitus, arterial sclerosis,multiple sclerosis, and other autoimmune reactions.

Accordingly, it is believed that it is advantageous to be able toselectively and controllably raise a target tissue temperature up to apredetermined temperature range over a short period of time, whilemaintaining the average temperature rise of the tissue at apredetermined temperature over a longer period of time. It is believedthat this induces the heat shock response in order to increase thenumber or activity of heat shock proteins in body tissue in response toinfection or other abnormalities. However, this must be done in acontrolled manner in order not to damage or destroy the tissue or thearea of the body being treated. It would also be desirable to maximizethe amount of heat shock protein activation within the cells of atargeted tissue during a single treatment session. The present inventionfulfills these needs, and provides other related advantages.

SUMMARY OF THE INVENTION

The present invention is directed to a process for heat treatingbiological tissues by applying treatment energy to a target tissue totherapeutically treat the target tissue. A first treatment to the targettissue is performed by generating treatment energy and repeatedlyapplying the treatment energy to the target tissue over a period of timeso as to controllably raise a temperature of the target tissue totherapeutically treat the target tissue without destroying orpermanently damaging the target tissue. The generated treatment energymay be pulsed or rapidly applied in succession. The target tissue maycomprise retinal tissue.

The energy parameters are selected so as to raise a target tissuetemperature up to 11° C. to achieve a therapeutic effect, wherein theaverage temperature rise of the tissue over several minutes ismaintained at or below a predetermined level so as not to permanentlydamage the target tissue. The energy parameters may be selected so thatthe target tissue temperature is raised between approximately 6° C. to11° C. at least during application of the energy to the target tissue.The average temperature rise of the target tissue over several minutesis maintained at 6° C. or less, such as at approximately 1° C. or lessover several minutes.

The treatment energy and application parameters are selected such so asto therapeutically treat the target tissue without destroying orpermanently damaging the target tissue. The selected energy andapplication parameters may comprise tissue application spot size orarea, average power or average power density, and exposure duration.Other parameters which may be selected include wavelength or frequencyand duty cycle. For example, the treatment energy and applicationparameters may be selected to have an average power density of 100-590watts per square centimeter of target tissue, a target tissueapplication spot size between 100-500 microns, and a train exposureduration of 500 milliseconds or less.

The treatment energy may comprise a light beam, a microwave, aradiofrequency or an ultrasound. A device may be inserted into a cavityof the body in order to apply the treatment energy to the tissue. Thetreatment energy may be applied to an exterior area of a body which isadjacent to the target tissue, or has a blood supply close to a surfaceof the exterior area of the body.

The treatment energy may comprise a radiofrequency between approximately3 to 6 megahertz (MHz). It may have a duty cycle of betweenapproximately 2.5% to 5%. It may have a pulsed train duration of betweenapproximately 0.2 to 0.4 seconds. The radiofrequency may be generatedwith a device having a coil radii of between approximately 2 and 6 mmand approximately 13 and 57 amp turns.

The treatment energy may comprise a microwave frequency of between 10 to20 gigahertz (GHz). The microwave may have a pulse train duration ofapproximately between 0.2 and 0.6 seconds. The microwave may have a dutycycle of between approximately 2% and 5%. The microwave may have anaverage power of between approximately 8 and 52 watts.

The treatment energy may comprise a pulsed light beam, such as one ormore laser light beams. The light beam may have a wavelength of betweenapproximately 570 nm to 1300 nm, and more preferably between 600 nm and1000 nm. The pulsed light beam may have a power of between approximately0.5 and 74 watts. The pulsed light beam has a duty cycle of less than10%, and preferably between 2.5% and 5%. The pulsed light beam may havea pulse train duration of approximately 0.1 and 0.6 seconds.

The treatment energy may comprise a pulsed ultrasound, having afrequency of between approximately 1 and 5 MHz. The ultrasound has atrain duration of approximately 0.1 and 05 seconds. The ultrasound mayhave a duty cycle of between approximately 2% and 10%. The ultrasoundhas a power of between approximately 0.46 and 28.6 watts.

The process of the present invention may comprise the steps of providinga plurality of energy emitters formed into an array. Treatment energy isgenerated from the plurality of emitters. The treatment energy isapplied to the target tissue, wherein the treatment energy has energyand application parameters selected so as to raise the target tissuetemperature sufficiently to create a therapeutic effect whilemaintaining an average temperature of the target tissue over severalminutes at or below a predetermined temperature so as not to destroy orpermanently damage the target tissue.

The first treatment comprises applying the treatment energy to thetarget tissue for a period of less than ten seconds, and more typicallyless than one second. The first treatment creates a level of heat shockprotein activation in the target tissue. The application of thetreatment energy to the target tissue is halted for an interval of timethat preferably exceeds the period of time of the first treatment. Theinterval of time may comprise several seconds to several minutes, suchas three seconds to three minutes, or preferably between ten seconds toninety seconds. After the interval of time and within a single treatmentsession, a second treatment is performed to the target tissue byrepeatedly reapplying the treatment energy to the target tissue so as tocontrollably raise the temperature of the target tissue totherapeutically treat the target tissue without destroying orpermanently damaging the target tissue. The second treatment increasesthe level of heat shock protein activation in the target tissue suchthat it is at a level which is higher than the level after the firsttreatment.

During an interval of time, typically comprising less than one second,between applications of treatment energy applied to a first area of thetarget tissue, the treatment energy may be applied to a second area ofthe target tissue sufficiently spaced apart from the first area of thetarget tissue to avoid thermal tissue damage of the target tissue. Thetreatment energy is repeatedly applied, in an alternating manner duringthe same treatment session, to each of the first and second areas of thetarget tissue until the predetermined number of energy applications toeach of the first and second areas of the target tissue has beenachieved.

When utilizing an array, a phase delay in the activation of the energyemitters of the array may be introduced to generate treatment energy ina phased manner using a predetermined delay of activation in order toapply treatment energy to each of the first and second areas of thetarget tissue. Alternatively, the energy emitters of the array may beactivated sequentially in order to apply treatment energy to each of thefirst and second areas of the target tissue.

Other features and advantages of the present invention will becomeapparent from the following more detailed description, taken inconjunction with the accompanying drawings, which illustrate, by way ofexample, the principles of the invention.

BRIEF DESCRIPTION OF THE DRAWINGS

The accompanying drawings illustrate the invention. In such drawings:

FIGS. 1A and 1B are graphs illustrating the average power of a lasersource compared to a source radius and pulse train duration of thelaser;

FIGS. 2A and 2B are graphs illustrating the time for the temperature todecay depending upon the laser source radius and wavelength;

FIGS. 3-6 are graphs illustrating the peak ampere turns for variousradiofrequencies, duty cycles, and coil radii;

FIG. 7 is a graph depicting the time for temperature rise to decaycompared to radiofrequency coil radius;

FIGS. 8 and 9 are graphs depicting the average microwave power comparedto microwave frequency and pulse train durations;

FIG. 10 is a graph depicting the time for the temperature to decay forvarious microwave frequencies;

FIG. 11 is a graph depicting the average ultrasound source powercompared to frequency and pulse train duration;

FIGS. 12 and 13 are graphs depicting the time for temperature decay forvarious ultrasound frequencies;

FIG. 14 is a graph depicting the volume of focal heated region comparedto ultrasound frequency;

FIG. 15 is a graph comparing equations for temperature over pulsedurations for an ultrasound energy source;

FIGS. 16 and 17 are graphs illustrating the magnitude of the logarithmof damage and HSP activation Arrhenius integrals as a function oftemperature and pulse duration;

FIG. 18 is a diagrammatic view of a light generating unit that producestimed series of pulses, having a light pipe extending therefrom, inaccordance with the present invention;

FIG. 19 is a cross-sectional view of a photostimulation delivery devicedelivering electromagnetic energy to target tissue, in accordance withthe present invention;

FIG. 20 is a diagrammatic view illustrating a system used to generate alaser light beam, in accordance with the present invention;

FIG. 21 is a diagrammatic view of optics used to generate a laser lightgeometric pattern, in accordance with the present invention;

FIG. 22 is a top plan view of an optical scanning mechanism, used inaccordance with the present invention;

FIG. 23 is a partially exploded view of the optical scanning mechanismof FIG. 22, illustrating the various component parts thereof;

FIG. 24 illustrates controlled offsets of exposure of an exemplarygeometric pattern grid of laser spots to treat the target tissue, inaccordance with an embodiment of the present invention;

FIG. 25 is a diagrammatic view illustrating the use of a geometricobject in the form of a line controllably scanned to treat an area ofthe target tissue;

FIG. 26 is a diagrammatic view similar to FIG. 25, but illustrating thegeometric line or bar rotated to treat the target tissue;

FIG. 27 is a diagrammatic view illustrating an alternate embodiment ofthe system used to generate laser light beams for treating tissue, inaccordance with the present invention;

FIG. 28 is a diagrammatic view illustrating yet another embodiment of asystem used to generate laser light beams to treat tissue in accordancewith the present invention;

FIG. 29 is a cross-sectional and diagrammatic view of an end of anendoscope inserted into the nasal cavity and treating tissue therein, inaccordance with the present invention;

FIG. 30 is a diagrammatic and partially cross-sectioned view of abronchoscope extending through the trachea and into the bronchus of alung and providing treatment thereto, in accordance with the presentinvention;

FIG. 31 is a diagrammatic view of a colonoscope providingphotostimulation to an intestinal or colon area of the body, inaccordance with the present invention;

FIG. 32 is a diagrammatic view of an endoscope inserted into a stomachand providing treatment thereto, in accordance with the presentinvention;

FIG. 33 is a partially sectioned perspective view of a capsuleendoscope, used in accordance with the present invention;

FIG. 34 is a diagrammatic view of a pulsed high intensity focusedultrasound for treating tissue internal the body, in accordance with thepresent invention;

FIG. 35 is a diagrammatic view for delivering therapy to the bloodstreamof a patient, through an earlobe, in accordance with the presentinvention;

FIG. 36 is a cross-sectional view of a stimulating therapy device of thepresent invention used in delivering photostimulation to the blood, viaan earlobe, in accordance with the present invention;

FIGS. 37A-37D are diagrammatic views illustrated in the application ofmicropulsed energy to different treatment areas during a predeterminedinterval of time, within a single treatment session, and reapplying theenergy to previously treated areas, in accordance with the presentinvention;

FIGS. 38-40 are graphs depicting the relationship of treatment power andtime in accordance with the embodiments of the present invention;

FIG. 41 is a graph depicting wavefront from two sources separated by adistance;

FIG. 42 is a depiction of a square array of square antennas or sources,which can be used in accordance with the present invention;

FIG. 43 is a graph depicting the shape of radiation pattern from asquare antenna array;

FIG. 44 is a graph depicting a form of typical radiation pattern alongan X-axis from a far field array;

FIG. 45 is a graph depicting an envelope of the pattern of FIG. 44;

FIG. 46 is another graph depicting the width of individual lines of thepattern of FIG. 44;

FIG. 47 is a plot graph depicting the determinant of the lineseparation;

FIG. 48 is a block diagram of components of a steerable array system;

FIG. 49 is a plot graph showing induced tissue temperature rise anddrops;

FIGS. 50-52 are graphs depicting variables of three different coils, inaccordance with the present invention;

FIG. 53 is a graph depicting the plots of FIGS. 50-52 superimposed uponone another;

FIG. 54 is a block diagram for an induction array which can be used inaccordance with the present invention;

FIGS. 55A and 55B are graphs depicting the behavior of HSP cellularsystem components over time following a sudden increase in temperature;

FIGS. 56A-56H are graphs depicting the behavior of HSP cellular systemcomponents in the first minute following a sudden increase intemperature;

FIGS. 57A and 57B are graphs illustrating variation in the activatedconcentrations of HSP and unactivated HSP in the cytoplasmic reservoirover an interval of one minute, in accordance with the presentinvention; and

FIG. 58 is a graph depicting the improvement ratios versus intervalbetween treatments, in accordance with the present invention.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

As shown in the accompanying drawings, and as more fully describedherein, the present invention is directed to a system and method fordelivering a pulsed energy, such as ultrasound, ultravioletradiofrequency, microwave radiofrequency, one or more light beams, andthe like, having energy parameters selected to cause a thermaltime-course in tissue to raise the tissue temperature over a shortperiod of time to a sufficient level to achieve a therapeutic effectwhile maintaining an average tissue temperature over a prolonged periodof time below a predetermined level so as to avoid permanent tissuedamage. It is believed that the creation of the thermal time-coursestimulates heat shock protein activation or production and facilitatesprotein repair without causing any damage.

The inventors have discovered that electromagnetic radiation can beapplied to retinal tissue in a manner that does not destroy or damagethe retinal tissue while achieving beneficial effects on eye diseases.More particularly, a laser light beam can be generated that istherapeutic, yet sublethal to retinal tissue cells and thus avoidsdamaging photocoagulation in the retinal tissue which providespreventative and protective treatment of the retinal tissue of the eye.It is believed that this may be due, at least in part, to thestimulation and activation of heat shock proteins and the facilitationof protein repair in the retinal tissue. This is disclosed in U.S.patent application Ser. No. 14/607,959 filed Jan. 28, 2015, Ser. No.13/798,523 filed Mar. 13, 2013, and Ser. No. 13/481,124 filed May 25,2012, the contents of which are hereby incorporated by reference as ifmade in full.

Various parameters of the light beam must be taken into account andselected so that the combination of the selected parameters achieve thetherapeutic effect while not permanently damaging the tissue. Theseparameters include laser wavelength, radius of the laser source ortissue application spot, laser power, total pulse train duration, andduty cycle of the pulse train.

The selection of these parameters may be determined by requiring thatthe Arrhenius integral for HSP activation be greater than 1 or unity.Arrhenius integrals are used for analyzing the impacts of actions onbiological tissue. See, for instance, The CRC Handbook of ThermalEngineering, ed. Frank Kreith, Springer Science and Business Media(2000). At the same time, the selected parameters must not permanentlydamage the tissue. Thus, the Arrhenius integral for damage may also beused, wherein the solved Arrhenius integral is less than 1 or unity.Alternatively, the FDA/FCC constraints on energy deposition per unitgram of tissue and temperature rise as measured over periods of minutesbe satisfied so as to avoid permanent tissue damage. The FDA/FCCrequirements on energy deposition and temperature rise are widely usedand can be referenced, for example, atwww.fda.gov/medicaldevices/deviceregulationandguidance/guidancedocuments/ucm073817.htm#attacha for electromagnetic sources, and Anastosio andP. LaRivero, ed., Emerging Imaging Technologies. CRC Press (2012), forultrasound sources. Generally speaking, tissue temperature rises ofbetween 6° C. and 11° C. can create therapeutic effect, such as byactivating heat shock proteins, whereas maintaining the average tissuetemperature over a prolonged period of time, such as over severalminutes, such as six minutes, below a predetermined temperature, such as6° C. and even 1° C. or less in certain circumstances, will notpermanently damage the tissue.

The inventors have discovered that generating a subthreshold, sublethalmicropulse laser light beam which has a wavelength greater than 532 nmand a duty cycle of less than 10% at a predetermined intensity or powerand a predetermined pulse length or exposure time creates desirableretinal photostimulation without any visible burn areas or tissuedestruction. More particularly, a laser light beam having a wavelengthof between 570 nm-1300 nm, and in a particularly preferred embodimentbetween 600 nm and 1100 nm, having a duty cycle of approximately2.5%-10% and a predetermined average power or power intensity (such asbetween 100-590 watts per square centimeter at the retina orapproximately 1 watt per laser spot for each treatment spot at theretina) and a predetermined pulse train length or exposure time (such asbetween 100 and 600 milliseconds or less) creates a sublethal, “truesubthreshold” retinal photostimulation in which all areas of the retinalpigment epithelium exposed to the laser irradiation are preserved andavailable to contribute therapeutically. In other words, the inventorshave found that raising the retinal tissue at least up to a therapeuticlevel but below a cellular or tissue lethal level recreates the benefitof the halo effect of the prior art methods without destroying, burningor otherwise damaging the retinal tissue. This is referred to herein assubthreshold diode micropulse laser treatment (SDM).

SDM does not produce laser-induced retinal damage (photocoagulation),and has no known adverse treatment effect, and has been reported to bean effective treatment in a number of retinal disorders (includingdiabetic macular edema (DME) proliferative diabetic retinopathy (PDR),macular edema due to branch retinal vein occlusion (BRVO), centralserous chorioretinopathy (CSR), reversal of drug tolerance, andprophylactic treatment of progressive degenerative retinopathies such asdry age-related macular degeneration, Stargardts' disease, conedystrophies, and retinitis pigmentosa. The safety of SDM is such that itmay be used transfoveally in eyes with 20/20 visual acuity to reduce therisk of visual loss due to early fovea-involving DME.

A mechanism through which SDM might work is the generation or activationof heat shock proteins (HSPs). Despite a near infinite variety ofpossible cellular abnormalities, cells of all types share a common andhighly conserved mechanism of repair: heat shock proteins (HSPs). HSPsare elicited almost immediately, in seconds to minutes, by almost anytype of cell stress or injury. In the absence of lethal cell injury,HSPs are extremely effective at repairing and returning the viable celltoward a more normal functional state. Although HSPs are transient,generally peaking in hours and persisting for a few days, their effectsmay be long lasting. HSPs reduce inflammation, a common factor in manydisorders.

Laser treatment can induce HSP production or activation and altercytokine expression. The more sudden and severe the non-lethal cellularstress (such as laser irradiation), the more rapid and robust HSPactivation. Thus, a burst of repetitive low temperature thermal spikesat a very steep rate of change (˜7° C. elevation with each 100 μsmicropulse, or 70,000° C./sec) produced by each SDM exposure isespecially effective in stimulating activation of HSPs, particularlycompared to non-lethal exposure to subthreshold treatment withcontinuous wave lasers, which can duplicate only the low average tissuetemperature rise.

Laser wavelengths below 550 nm produce increasingly cytotoxicphotochemical effects. The lower wavelength limit realistically usableby the process of the present invention is determined by the undesirableabsorption by the visual pigments and other absorbers, including blood,the lens of the eye, etc. At approximately 570 nm, the sum of theoptical densities of the long wavelength sensitive and medium wavelengthsensitive visual pigments in the eye and the blood exceeds the opticaldensity of the melanin. The absorption is dominated by melanin between570 nm and 650 nm, where above 650 nm the absorption is practically alldue to the melanin in the RPE. However, at higher wavelengths, such asabove 1300 nm, there is a decrease in melanin absorption with increasingabsorption by the water in the vitreous of the eye. At 1300 nm, forinstance, the melanin absorbance is only 0.048 of what it is at 810 nm,and the radiation power due to this effect alone would have to beincreased by a factor of 20 compared to the power at 810 nm to achievethe same temperature increase. Accordingly, the present invention can beperformed at a broad range of wavelengths between 570 nm to 1300 nm,with the more preferable range of wavelengths being 600 nm to 1100 nm,and an even more preferable range of wavelengths of 650 nm to 900 nm,with the particularly preferred operating wavelength at approximately810 nm. At these wavelengths, the melanin absorption is dominant and theheating primarily in the desired RPE and the wavelength is at a safedistance from the wavelengths where appreciable absorption occurs in thevisual pigments as shorter wavelengths or water at longer wavelengths,which will create undesirable heating of the eye and other tissues. At810 nm, SDM produces photothermal, rather than photochemical, cellularstress. Thus, SDM is able to affect the tissue without damaging it.

It has been found that the average required treatment power betweentissue reset and tissue damage can be calculated with the wavelengthused, the radiation train duration, preferably being between 0.03 and0.8 seconds and a retinal application spot by the radiation beingbetween 10 and 500 microns. A duty cycle of less than 10% and preferablybetween 2.5% and 5% with a total pulse duration of between 100milliseconds and 600 milliseconds has been found to be effective. Thecorresponding peak powers, during the individual pulse, are obtainedfrom the average powers by dividing by the duty cycle. The average powercan vary between 0.0000069 to 37.5 watts within a wavelength between 570nm-1300 nm, a pulse train duration between 30-800 milliseconds, and atreatment spot between 10-700 microns.

The clinical benefits of SDM are thus primarily produced by sub-morbidphotothermal cellular HSP activation. In dysfunctional cells, HSPstimulation by SDM results in normalized cytokine expression, andconsequently improved structure and function. The therapeutic effects ofthis “low-intensity” laser/tissue interaction are then amplified by“high-density” laser application, recruiting all the dysfunctional cellsin the targeted tissue area by densely/confluently treating a largetissue area, including all areas of pathology, thereby maximizing thetreatment effect. These principles define the treatment strategy of SDMdescribed herein.

Because normally functioning cells are not in need of repair, HSPstimulation in normal cells would tend to have no notable clinicaleffect. The “patho-selectivity” of near infrared laser effects, such asSDM, affecting sick cells but not affecting normal ones, on various celltypes is consistent with clinical observations of SDM. SDM has beenreported to have a clinically broad therapeutic range, unique amongretinal laser modalities, consistent with American National StandardsInstitute “Maximum Permissible Exposure” predictions. While SDM maycause direct photothermal effects such as entropic protein unfolding anddisaggregation, SDM appears optimized for clinically safe and effectivestimulation of HSP-mediated repair.

As noted above, while SDM stimulation of HSPs is non-specific withregard to the disease process, the result of HSP mediated repair is byits nature specific to the state of the dysfunction. HSPs tend to fixwhat is wrong, whatever that might be. Thus, the observed effectivenessof SDM in retinal conditions as widely disparate as BRVO, DME, PDR, CSR,age-related and genetic retinopathies, and drug-tolerant NAMD.Conceptually, this facility can be considered a sort of “Reset toDefault” mode of SDM action. For the wide range of disorders in whichcellular function is critical, SDM normalizes cellular function bytriggering a “reset” (to the “factory default settings”) viaHSP-mediated cellular repair.

The inventors have found that SDM treatment of patients suffering fromage-related macular degeneration (AMD) can slow the progress or evenstop the progression of AMD. Most of the patients have seen significantimprovement in dynamic functional log MAR mesoptic visual acuity andmesoptic contrast visual acuity after the SDM treatment. It is believedthat SDM works by targeting, preserving, and “normalizing” (movingtoward normal) function of the retinal pigment epithelium (RPE).

SDM has also been shown to stop or reverse the manifestations of thediabetic retinopathy disease state without treatment-associated damageor adverse effects, despite the persistence of systemic diabetesmellitus. On this basis it is hypothesized that SDM might work byinducing a return to more normal cell function and cytokine expressionin diabetes-affected RPE cells, analogous to hitting the “reset” buttonof an electronic device to restore the factory default settings. Basedon the above information and studies, SDM treatment may directly affectcytokine expression via heat shock protein (HSP) activation in thetargeted tissue.

As heat shock proteins play a role in responding to a large number ofabnormal conditions in body tissue other than eye tissue, it is believedthat similar systems and methodologies can be advantageously used intreating such abnormal conditions, infections, etc. As such, the presentinvention is directed to the controlled application of ultrasound orelectromagnetic radiation to treat abnormal conditions includinginflammations, autoimmune conditions, and cancers that are accessible bymeans of fiber optics of endoscopes or surface probes as well as focusedelectromagnetic/sound waves. For example, cancers on the surface of theprostate that have the largest threat of metastasizing can be accessedby means of fiber optics in a proctoscope. Colon tumors can be accessedby an optical fiber system, like those used in colonoscopy.

As indicated above, subthreshold diode micropulse laser (SDM)photostimulation has been effective in stimulating direct repair ofslightly misfolded proteins in eye tissue. Besides HSP activation,another way this may occur is because the spikes in temperature causedby the micropulses in the form of a thermal time-course allows diffusionof water inside proteins, and this allows breakage of thepeptide-peptide hydrogen bonds that prevent the protein from returningto its native state. The diffusion of water into proteins results in anincrease in the number of restraining hydrogen bonds by a factor on theorder of a thousand. Thus, it is believed that this process could beapplied to other tissues and diseases advantageously as well.

As explained above, the energy source to be applied to the target tissuewill have energy and operating parameters which must be determined andselected so as to achieve the therapeutic effect while not permanentlydamaging the tissue. Using a light beam energy source, such as a laserlight beam, as an example, the laser wavelength, duty cycle and totalpulse train duration parameters must be taken into account. Otherparameters which can be considered include the radius of the lasersource as well as the average laser power. Adjusting or selecting one ofthese parameters can have an effect on at least one other parameter.

FIGS. 1A and 1B illustrate graphs showing the average power in watts ascompared to the laser source radius (between 0.1 cm and 0.4 cm) andpulse train duration (between 0.1 and 0.6 seconds). FIG. 1A shows awavelength of 880 nm, whereas FIG. 1B has a wavelength of 1000 nm. Itcan be seen in these figures that the required power decreasesmonotonically as the radius of the source decreases, as the total trainduration increases, and as the wavelength decreases. The preferredparameters for the radius of the laser source is 1 mm-4 mm. For awavelength of 880 nm, the minimum value of power is 0.55 watts, with aradius of the laser source being 1 mm, and the total pulse trainduration being 600 milliseconds. The maximum value of power for the 880nm wavelength is 52.6 watts when the laser source radius is 4 mm and thetotal pulse drain duration is 100 milliseconds. However, when selectinga laser having a wavelength of 1000 nm, the minimum power value is 0.77watts with a laser source radius of 1 mm and a total pulse trainduration of 600 milliseconds, and a maximum power value of 73.6 wattswhen the laser source radius is 4 mm and the total pulse duration is 100milliseconds. The corresponding peak powers, during an individual pulse,are obtained from the average powers by dividing by the duty cycle.

The volume of the tissue region to be heated is determined by thewavelength, the absorption length in the relevant tissue, and by thebeam width. The total pulse duration and the average laser powerdetermine the total energy delivered to heat up the tissue, and the dutycycle of the pulse train gives the associated spike, or peak, powerassociated with the average laser power. Preferably, the pulsed energysource energy parameters are selected so that approximately 20 to 40joules of energy is absorbed by each cubic centimeter of the targettissue.

The absorption length is very small in the thin melanin layer in theretinal pigmented epithelium. In other parts of the body, the absorptionlength is not generally that small. In wavelengths ranging from 400 nmto 2000 nm, the penetration depth and skin is in the range of 0.5 mm to3.5 mm. The penetration depth into human mucous tissues is in the rangeof 0.5 mm to 6.8 mm. Accordingly, the heated volume will be limited tothe exterior or interior surface where the radiation source is placed,with a depth equal to the penetration depth, and a transverse dimensionequal to the transverse dimension of the radiation source. Since thelight beam energy source is used to treat diseased tissues near externalsurfaces or near internal accessible surfaces, a source radii of between1 mm to 4 mm and operating a wavelength of 880 nm yields a penetrationdepth of approximately 2.5 mm and a wavelength of 1000 nm yields apenetration depth of approximately 3.5 mm.

It has been determined that the target tissue can be heated to up toapproximately 11° C. for a short period of time, such as less than onesecond, to create the therapeutic effect of the invention whilemaintaining the target tissue average temperature to a lower temperaturerange, such as less than 6° C. or even 1° C. or less over a prolongedperiod of time, such as several minutes. The selection of the duty cycleand the total pulse train duration provide time intervals in which theheat can dissipate. A duty cycle of less than 10%, and preferablybetween 2.5% and 5%, with a total pulse duration of between 100milliseconds and 600 milliseconds has been found to be effective. FIGS.2A and 2B illustrate the time to decay from 10° C. to 1° C. for a lasersource having a radius of between 0.1 cm and 0.4 cm with the wavelengthbeing 880 nm in FIG. 2A and 1000 nm in FIG. 2B. It can be seen that thetime to decay is less when using a wavelength of 880 nm, but eitherwavelength falls within the acceptable requirements and operatingparameters to achieve the benefits of the present invention while notcausing permanent tissue damage.

It has been found that the average temperature rise of the desiredtarget region increasing at least 6° C. and up to 11° C., and preferablyapproximately 10° C., during the total irradiation period results in HSPactivation. The control of the target tissue temperature is determinedby choosing source and target parameters such that the Arrheniusintegral for HSP activation is larger than 1, while at the same timeassuring compliance with the conservative FDA/FCC requirements foravoiding damage or a damage Arrhenius integral being less than 1.

In order to meet the conservative FDA/FCC constraints to avoid permanenttissue damage, for light beams and other electromagnetic radiationsources, the average temperature rise of the target tissue over anysix-minute period is 1° C. or less. FIGS. 2A and 2B above illustrate thetypical decay times required for the temperature in the heated targetregion to decrease by thermal diffusion from a temperature rise ofapproximately 10° C. to 1° C. as can be seen in FIG. 2A when thewavelength is 880 nm and the source diameter is 1 millimeter, thetemperature decay time is 16 seconds. The temperature decay time is 107seconds when the source diameter is 4 mm. As shown in FIG. 2B, when thewavelength is 1000 nm, the temperature decay time is 18 seconds when thesource diameter is 1 mm and 136 seconds when the source diameter is 4mm. This is well within the time of the average temperature rise beingmaintained over the course of several minutes, such as 6 minutes orless. While the target tissue's temperature is raised, such as toapproximately 10° C., very quickly, such as in a fraction of a secondduring the application of the energy source to the tissue, therelatively low duty cycle provides relatively long periods of timebetween the pulses of energy applied to the tissue and the relativelyshort pulse train duration ensure sufficient temperature diffusion anddecay within a relatively short period of time comprising severalminutes, such as 6 minutes or less, that there is no permanent tissuedamage.

The parameters differ for the individual energy sources, includingmicrowave, infrared lasers, radiofrequency and ultrasound, because theabsorption properties of tissues differ for these different types ofenergy sources. The tissue water content can vary from one tissue typeto another, however, there is an observed uniformity of the propertiesof tissues at normal or near normal conditions which has allowedpublication of tissue parameters that are widely used by clinicians indesigning treatments. Below are tables illustrating the properties ofelectromagnetic waves in biological media, with Table 1 relating tomuscle, skin and tissues with high water content, and Table 2 relatingto fat, bone and tissues with low water content.

TABLE 1 Properties of Electromagnetic Waves in Biological Media: Muscle,Skin, and Tissues with High Water Content Reflection CoefficientWavelength Dielectric Conductivity Wavelength Depth of Air-MuscleMuscle-Fat Frequency in Air Constant σH λH Penetration InterfaceInterface (MHz) (cm) €H (mho/m) (cm) (cm) r ø r ø 1 30000 2000 0.400 43691.3 0.982 +179 10 3000 160 0.625 118 21.6 0.956 +178 27.12 1106 1130.612 68.1 14.3 0.925 +177 0.651 −11.13 40.68 738 97.3 0.693 51.3 11.20.913 +176 0.652 −10.21 100 300 71.7 0.889 27 6.66 0.881 +175 0.650−7.96 200 150 56.5 1.28 16.6 4.79 0.844 +175 0.612 −8.06 300 100 54 1.3711.9 3.89 0.825 +175 0.592 −8.14 433 69.3 53 1.43 8.76 3.57 0.803 +1750.562 −7.06 750 40 52 1.54 5.34 3.18 0.779 +176 0.532 −5.69 915 32.8 511.60 4.46 3.04 0.772 +177 0.519 −4.32 1500 20 49 1.77 2.81 2.42 0.761+177 0.506 −3.66 2450 12.2 47 2.21 1.76 1.70 0.754 +177 0.500 −3.88 300010 46 2.26 1.45 1.61 0.751 +178 0.495 −3.20 5000 6 44 3.92 0.89 0.7880.749 +177 0.502 −4.95 5800 5.17 43.3 4.73 0.775 0.720 0.746 +177 0.502−4.29 8000 3.75 40 7.65 0.578 0.413 0.744 +176 0.513 −6.65 10000 3 39.910.3 0.464 0.343 0.743 +176 0.518 −5.95

TABLE 2 Properties of Electromagnetic Waves in Biological Media: Fat,Bone, and Tissues with Low Water Content Reflection CoefficientWavelength Dielectric Conductivity Wavelength Depth of Air-FatFat-Muscle Frequency in Air Constant σL, λL Penetration InterfaceInterface (MHz) (cm) €L (mmho/m) (cm (cm) r ø r ø 1 30000 10 3000 27.121106 20 10.9-43.2 241 159 0.660 +174 0.651 +169 40.68 738 14.6 12.6-52.8187 118 0.617 +173 0.652 +170 100 300 7.45 19.1-75.9 106 60.4 0.511 +1680.650 +172 200 150 5.95 25.8-94.2 59.7 39.2 0.458 +168 0.612 +172 300100 5.7 31.6-107  41 32.1 0.438 +169 0.592 +172 433 69.3 5.6 37.9-118 28.8 26.2 0.427 +170 0.562 +173 750 40 5.6 49.8-138  16.8 23 0.415 +1730.532 +174 915 32.8 5.6 55.6-147  13.7 17.7 0.417 +173 0.519 +176 150020 5.6 70.8-171  8.41 13.9 0.412 +174 0.506 +176 2450 12.2 5.5 96.4-213 5.21 11.2 0.406 +176 0.500 +176 3000 10 5.5 110-234 4.25 9.74 0.406 +1760.495 +177 5000 6 5.5 162-309 2.63 6.67 0.393 +176 0.502 +175 5900 5.175.05 186-338 2.29 5.24 0.388 +176 0.502 +176 8000 3.75 4.7 255-431 1.734.61 0.371 +176 0.513 +173 - 10000 3 4.5 324-549 1.41 3.39 0.363 +1750.518 +174, -

The absorption lengths of radiofrequency in body tissue are longcompared to body dimensions. Consequently, the heated region isdetermined by the dimensions of the coil that is the source of theradiofrequency energy rather than by absorption lengths. Long distancesr from a coil the magnetic (near) field from a coil drops off as 1/r³.At smaller distances, the electric and magnetic fields can be expressedin terms of the vector magnetic potential, which in turn can beexpressed in closed form in terms of elliptic integrals of the first andsecond kind. The heating occurs only in a region that is comparable insize to the dimensions of the coil source itself. Accordingly, if it isdesired to preferentially heat a region characterized by a radius, thesource coil will be chosen to have a similar radius. The heating dropsoff very rapidly outside of a hemispherical region of radius because ofthe 1/r³ drop off of the magnetic field. Since it is proposed to use theradiofrequency the diseased tissue accessible only externally or frominner cavities, it is reasonable to consider a coil radii of betweenapproximately 2 to 6 mm.

The radius of the source coil(s) as well as the number of ampere turns(NI) in the source coils give the magnitude and spatial extent of themagnetic field, and the radiofrequency is a factor that relates themagnitude of the electric field to the magnitude of the magnetic field.The heating is proportional to the product of the conductivity and thesquare of the electric field. For target tissues of interest that arenear external or internal surfaces, the conductivity is that of skin andmucous tissue. The duty cycle of the pulse train as well as the totaltrain duration of a pulse train are factors which affect how much totalenergy is delivered to the tissue.

Preferred parameters for a radiofrequency energy source have beendetermined to be a coil radii between 2 and 6 mm, radiofrequencies inthe range of 3-6 MHz, total pulse train durations of 0.2 to 0.4 seconds,and a duty cycle of between 2.5% and 5%. FIGS. 3-6 show how the numberof ampere turns varies as these parameters are varied in order to give atemperature rise that produces an Arrhenius integral of approximatelyone or unity for HSP activation. With reference to FIG. 3, for an RFfrequency of 6 MHz, a pulse train duration of between 0.2 and 0.4seconds, a coil radius between 0.2 and 0.6 cm, and a duty cycle of 5%,the peak ampere turns (NI) is 13 at the 0.6 cm coil radius and 20 at the0.2 cm coil radius. For a 3 MHz frequency, as illustrated in FIG. 4, thepeak ampere turns is 26 when the pulse train duration is 0.4 seconds andthe coil radius is 0.6 cm and the duty cycle is 5%. However, with thesame 5% duty cycle, the peak ampere turns is 40 when the coil radius is0.2 cm and the pulse train duration is 0.2 seconds. A duty cycle of 2.5%is used in FIGS. 5 and 6. This yields, as illustrated in FIG. 5, 18 ampturns for a 6 MHz radiofrequency having a coil radius of 0.6 cm and apulse train duration of 0.4 seconds, and 29 amp turns when the coilradius is only 0.2 cm and the pulse train duration is 0.2 seconds. Withreference to FIG. 6, with a duty cycle of 2.5% and a radiofrequency of 3MHz, the peak ampere turns is 36 when the pulse train duration is 0.4seconds and the coil radius is 0.6 cm, and 57 amp turns when the pulsetrain duration is 0.2 seconds and the coil radius is 0.2 cm.

The time, in seconds, for the temperature rise to decay fromapproximately 10° C. to approximately 1° C. for coil radii between 0.2cm and 0.6 cm is illustrated for a radiofrequency energy source in FIG.7. The temperature decay time is approximately 37 seconds when theradiofrequency coil radius is 0.2 cm, and approximately 233 seconds whenthe radiofrequency coil radius is 0.5 cm. When the radiofrequency coilradius is 0.6 cm, the decay time is approximately 336 seconds, which isstill within the acceptable range of decay time, but at an upper rangethereof.

Microwaves are another electromagnetic energy source which can beutilized in accordance with the present invention. The frequency of themicrowave determines the tissue penetration distance. The gain of aconical microwave horn is large compared to the microwave wavelength,indicating under those circumstances that the energy is radiated mostlyin a narrow forward load. Typically, a microwave source used inaccordance with the present invention has a linear dimension on theorder of a centimeter or less, thus the source is smaller than thewavelength, in which case the microwave source can be approximated as adipole antenna. Such small microwave sources are easier to insert intointernal body cavities and can also be used to radiate externalsurfaces. In that case, the heated region can be approximated by ahemisphere with a radius equal to the absorption length of the microwavein the body tissue being treated. As the microwaves are used to treattissue near external surfaces or surfaces accessible from internalcavities, frequencies in the 10-20 GHz range are used, wherein thecorresponding penetration distances are only between approximately 2 and4 mm.

The temperature rise of the tissue using a microwave energy source isdetermined by the average power of the microwave and the total pulsetrain duration. The duty cycle of the pulse train determines the peakpower in a single pulse in a train of pulses. As the radius of thesource is taken to be less than approximately 1 centimeter, andfrequencies between 10 and 20 GHz are typically used, a resulting pulsetrain duration of 0.2 and 0.6 seconds is preferred.

The required power decreases monotonically as the train durationincreases and as the microwave frequency increases. For a frequency of10 GHz, the average power is 18 watts when the pulse train duration is0.6 seconds, and 52 watts when the pulse train duration is 0.2 seconds.For a 20 GHz microwave frequency, an average power of 8 watts is usedwhen the pulse train is 0.6 seconds, and can be 26 watts when the pulsetrain duration is only 0.2 seconds. The corresponding peak power areobtained from the average power simply by dividing by the duty cycle.

With reference now to FIG. 8, a graph depicts the average microwavepower in watts of a microwave having a frequency of 10 GHz and a pulsetrain duration from between 0.2 seconds and 0.6 seconds. FIG. 9 is asimilar graph, but showing the average microwave power for a microwavehaving a frequency of 20 GHz. Thus, it will be seen that the averagemicrowave source power varies as the total train duration and microwavefrequency vary. The governing condition, however, is that the Arrheniusintegral for HSP activation in the heated region is approximately 1.

With reference to FIG. 10, a graph illustrates the time, in seconds, forthe temperature to decay from approximately 10° C. to 1° C. compared tomicrowave frequencies between 58 MHz and 20000 MHz. The minimum andmaximum temperature decay for the preferred range of microwavefrequencies are 8 seconds when the microwave frequency is 20 GHz, and 16seconds when the microwave frequency is 10 GHz.

Utilizing ultrasound as an energy source enables heating of surfacetissue, and tissues of varying depths in the body, including rather deeptissue. The absorption length of ultrasound in the body is rather long,as evidenced by its widespread use for imaging. Accordingly, ultrasoundcan be focused on target regions deep within the body, with the heatingof a focused ultrasound beam concentrated mainly in the approximatelycylindrical focal region of the beam. The heated region has a volumedetermined by the focal waist of the airy disc and the length of thefocal waist region, that is the confocal parameter. Multiple beams fromsources at different angles can also be used, the heating occurring atthe overlapping focal regions.

For ultrasound, the relevant parameters for determining tissuetemperature are frequency of the ultrasound, total train duration, andtransducer power when the focal length and diameter of the ultrasoundtransducer is given. The frequency, focal length, and diameter determinethe volume of the focal region where the ultrasound energy isconcentrated. It is the focal volume that comprises the target volume oftissue for treatment. Transducers having a diameter of approximately 5cm and having a focal length of approximately 10 cm are readilyavailable. Favorable focal dimensions are achieved when the ultrasoundfrequency is between 1 and 5 MHz, and the total train duration is 0.1 to0.5 seconds. For example, for a focal length of 10 cm and the transducerdiameter of 5 cm, the focal volumes are 0.02 cc at 5 MHz and 2.36 cc at1 MHz.

With reference now to FIG. 11, a graph illustrates the average sourcepower in watts compared to the frequency (between 1 MHz and 5 MHz), andthe pulse train duration (between 0.1 and 0.5 seconds). A transducerfocal length of 10 cm and a source diameter of 5 cm have been assumed.The required power to give the Arrhenius integral for HSP activation ofapproximately 1 decreases monotonically as the frequency increases andas the total train duration increases. Given the preferred parameters,the minimum power for a frequency of 1 GHz and a pulse train duration of0.5 seconds is 5.72 watts, whereas for the 1 GHz frequency and a pulsetrain duration of 0.1 seconds the maximum power is 28.6 watts. For a 5GHz frequency, 0.046 watts is required for a pulse train duration of 0.5seconds, wherein 0.23 watts is required for a pulse train duration of0.1 seconds. The corresponding peak power during an individual pulse isobtained simply by dividing by the duty cycle.

FIG. 12 illustrates the time, in seconds, for the temperature to diffuseor decay from 10° C. to 6° C. when the ultrasound frequency is between 1and 5 MHz. FIG. 13 illustrates the time, in seconds, to decay fromapproximately 10° C. to approximately 1° C. for ultrasound frequenciesfrom 1 to 5 MHz. For the preferred focal length of 10 cm and thetransducer diameter of 5 cm, the maximum time for temperature decay is366 seconds when the ultrasound frequency is 1 MHz, and the minimumtemperature decay is 15 seconds when the microwave frequency is 5 MHz.As the FDA only requires the temperature rise be less than 6° C. fortest times of minutes, the 366 second decay time at 1 MHz to get to arise of 1° C. over the several minutes is allowable. As can be seen inFIGS. 12 and 13, the decay times to a rise of 6° C. are much smaller, bya factor of approximately 70, than that of 1° C.

FIG. 14 illustrates the volume of focal heated region, in cubiccentimeters, as compared to ultrasound frequencies from between 1 and 5MHz. Considering ultrasound frequencies in the range of 1 to 5 MHz, thecorresponding focal sizes for these frequencies range from 3.7 mm to 0.6mm, and the length of the focal region ranges from 5.6 cm to 1.2 cm. Thecorresponding treatment volumes range from between approximately 2.4 ccand 0.02 cc.

Examples of parameters giving a desired HSP activation Arrheniusintegral greater than 1 and damage Arrhenius integral less than 1 is atotal ultrasound power between 5.8-17 watts, a pulse duration of 0.5seconds, an interval between pulses of 5 seconds, with total number ofpulses 10 within the total pulse stream time of 50 seconds. The targettreatment volume would be approximately 1 mm on a side. Larger treatmentvolumes could be treatable by an ultrasound system similar to a laserdiffracted optical system, by applying ultrasound in multiplesimultaneously applied adjacent but separated and spaced columns. Themultiple focused ultrasound beams converge on a very small treatmenttarget within the body, the convergence allowing for a minimal heatingexcept at the overlapping beams at the target. This area would be heatedand stimulate the activation of HSPs and facilitate protein repair bytransient high temperature spikes. However, given the pulsating aspectof the invention as well as the relatively small area being treated atany given time, the treatment is in compliance with FDA/FCC requirementsfor long term (minutes) average temperature rise <1K. An importantdistinction of the invention from existing therapeutic heatingtreatments for pain and muscle strain is that there are no high T spikesin existing techniques, and these are required for efficientlyactivating HSPs and facilitating protein repair to provide healing atthe cellular level.

The pulse train mode of energy delivery has a distinct advantage over asingle pulse or gradual mode of energy delivery, as far as theactivation of remedial HSPs and the facilitation of protein repair isconcerned. There are two considerations that enter into this advantage:

First, a big advantage for HSP activation and protein repair in an SDMenergy delivery mode comes from producing a spike temperature of theorder of 10° C. This large rise in temperature has a big impact on theArrhenius integrals that describe quantitatively the number of HSPs thatare activated and the rate of water diffusion into the proteins thatfacilitates protein repair. This is because the temperature enters intoan exponential that has a big amplification effect.

It is important that the temperature rise not remain at the high value(10° C. or more) for long, because then it would violate the FDA and FCCrequirements that over periods of minutes the average temperature risemust be less than 1° C. (or in the case of ultrasound 6° C.).

An SDM mode of energy delivery uniquely satisfies both of theseforegoing considerations by judicious choice of the power, pulse time,pulse interval, and the volume of the target region to be treated. Thevolume of the treatment region enters because the temperature must decayfrom its high value of the order of 10° C. fairly rapidly in order forthe long term average temperature rise not to exceed the long termFDA/FCC limit of 6° C. for ultrasound frequencies and 1° C. or less forelectromagnetic radiation energy sources.

For a region of linear dimension L, the time that it takes the peaktemperature to e-fold in tissue is roughly L²/16 D, where D=0.00143cm²/sec is the typical heat diffusion coefficient. For example, if L=1mm, the decay time is roughly 0.4 sec. Accordingly, for a region 1 mm ona side, a train consisting of 10 pulses each of duration 0.5 seconds,with an interval between pulses of 5 second can achieve the desiredmomentary high rise in temperature while still not exceeding an averagelong term temperature rise of 1° C. This is demonstrated further below.

The limitation of heated volume is the reason why RF electromagneticradiation is not as good of a choice for SDM-type treatment of regionsdeep with the body as ultrasound. The long skin depths (penetrationdistances) and Ohmic heating all along the skin depth results in a largeheated volume whose thermal inertia does not allow both the attainmentof a high spike temperature that activates HSPs and facilitates proteinrepair, and the rapid temperature decay that satisfies the long term FDAand FCC limit on average temperature rise.

Ultrasound has already been used to therapeutically heat regions of thebody to ease pain and muscle strain. However, the heating has notfollowed the SDM-type protocol and does not have the temperature spikesthat are responsible for the excitation of HSPs.

Consider, then, a group of focused ultrasound beams that are directed ata target region deep within the body. To simplify the mathematics,suppose that the beams are replaced by a single source with a sphericalsurface shape that is focused on the center of the sphere. Theabsorption lengths of ultrasound can be fairly long. Table 3 below showstypical absorption coefficients for ultrasound at 1 MHz. The absorptioncoefficients are roughly proportional to the frequency.

TABLE 3 Typical absorption coefficients for 1 MHz ultrasound in bodytissue: Body Tissue Attenuation Coefficient at 1 MHz (cm⁻¹) Water0.00046 Blood 0.0415 Fat 0.145 Liver 0.115-0.217 Kidney 0.23 Muscle 0.3-0.76 Bone 1.15

Assuming that the geometric variation of the incoming radiation due tothe focusing dominates any variation due to attenuation, the intensityof the incoming ultrasound at a distance r from the focus can be writtenapproximately as:

I(r)=P/(4πr ²)  [1]

where P denotes the total ultrasound power.

The temperature rise at the end of a short pulse of duration t_(p) at ris then

dT(t _(p))=Pαt _(p)/(4πC _(v) r ²)  [2]

where α is the absorption coefficient and C_(v) is the specific volumeheat capacity. This will be the case until the r is reached at which theheat diffusion length at t_(p) becomes comparable to r, or thediffraction limit of the focused beam is reached. For smaller r, thetemperature rise is essentially independent of r. As an example, supposethe diffraction limit is reached at a radial distance that is smallerthan that determined by heat diffusion. Then

r _(dif)=(4Dt _(p))^(1/2)  [3]

where D is the heat diffusion coefficient, and for r<r_(dif), thetemperature rise at t_(p) is

dT(r _(dif) ,t _(p))=3Pα/(8πC _(v) D) when r<r _(dif)  [4]

Thus, at the end of the pulse, we can write for the temperature rise:

dT _(p)(r)={Pαt _(p)/(4πC _(v)}[(6/r _(dif) ²)U{r _(dif) −r)+(1/r²)U(r−r _(dif))]  [5]

On applying the Green's function for the heat diffusion equation,

G(r,t)=(4ΩDt)^(−3/2)exp[−r ²/(4Dt)]  [6]

to this initial temperature distribution, we find that the temperaturedT(t) at the focal point r=0 at a time t is

dT(t)=[dT _(o)/{(½)±(π^(1/2)/6)}][(½)(t _(p) /t)^(3/2)+(π^(1/2)/6)(t_(p) /t)]  [7]

with

dT _(o)=3Pα/(8πC _(v) D)  [8]

A good approximation to eq. [7] is provided by:

dT(t)≈dT _(o)(t _(p) /t)^(3/2)  [9]

as can be seen in FIG. 15, which is a comparison of eqs. [7] and [9] fordT(t)/dT_(o) at the target treatment zone. The bottom curve is theapproximate expression of eq [9].The Arrhenius integral for a train of N pulses can now be evaluated withthe temperature rise given by eq. [9]. In this expression,

dT _(N)(t)=ΣdT(t−nt _(l))  [11]

where dT(t−nt_(I)) is the expression of eq. [9] with t replaced byt-nt_(l) and with t_(l) designating the interval between pulses.

The Arrhenius integral can be evaluated approximately by dividing theintegration interval into the portion where the temperature spikes occurand the portion where the temperature spike is absent. The summationover the temperature spike contribution can be simplified by applyingLaplace's end point formula to the integral over the temperature spike.In addition, the integral over the portion when the spikes are absentcan be simplified by noting that the non-spike temperature rise veryrapidly reaches an asymptotic value, so that a good approximation isobtained by replacing the varying time rise by its asymptotic value.When these approximations are made, eq. [10] becomes:

Ω=AN[{t _(p)(2k _(B) T _(o) ²/(3EdT _(o))}exp[−(E/k _(B))1/(T _(o) +dT_(o) +dT _(N)(Nt _(l)))]+exp[−(E/k _(B))1/(T _(o) +dT _(N)(Nt_(l)))]]  [12]

where

dT _(N)(Nt _(l))≈2.5dT _(o)(t _(p) /t _(l))^(3/2)  [13]

(The 2.5 in eq. [13] arises from the summation over n of (N−n)^(−3/2)and is the magnitude of the harmonic number (N,3/2) for typical N ofinterest.)

It is interesting to compare this expression with that for SDM appliedto the retina. The first term is very similar to that from the spikecontribution in the retina case, except that the effective spikeinterval is reduced by a factor of 3 for this 3D converging beam case.The second term, involving dT_(N)(Nt_(l)) is much smaller than in theretina case. There the background temperature rise was comparable inmagnitude to the spike temperature rise. But here in the converging beamcase, the background temperature rise is much smaller by the ratio(t_(p)/t_(l))^(3/2). This points up the importance of the spikecontribution to the activation or production of HSP's and thefacilitation of protein repair, as the background temperature rise whichis similar to the rise in a continuous ultrasound heating case isinsignificant compared to the spike contribution. At the end of thepulse train, even this low background temperature rise rapidlydisappears by heat diffusion.

FIGS. 16 and 17 show the magnitude of the logarithm of the Arrheniusintegrals for damage and for HSP activation or production as a functionof dT_(o) for a pulse duration t_(p)=0.5 sec, pulse interval t_(l)=10sec, and total number of pulses N=10. Logarithm of Arrhenius integrals[eq. 12] for damage and for HSP activation as a function of thetemperature rise in degrees Kelvin from a single pulse dT_(o), for apulse duration t_(o)=0.5 sec., pulse interval t_(l)=10 sec., and a totalnumber of ultrasound pulses N=10. FIG. 16 shows the logarithm of thedamage integral with the Arrhenius constants A=8.71×10³³ sec⁻¹ andE=3.55×10⁻¹² ergs. FIG. 17 shows the logarithm of the HSP activationintegral with the Arrhenius constants A=1.24×10²⁷ sec⁻¹ and E=2.66×10⁻¹²ergs. The graphs in FIGS. 16 and 17 show that Ω_(damage) does not exceed1 until dT_(o) exceeds 11.3 K, whereas Ω_(hsp) is greater than 1 overthe whole interval shown, the desired condition for cellular repairwithout damage.

Equation [8] shows that when α=0.1 cm⁻¹, a dT_(o) of 11.5 K can beachieved with a total ultrasound power of 5.8 watts. This is easilyachievable. If α is increased by a factor of 2 or 3, the resulting poweris still easily achievable. The volume of the region where thetemperature rise is constant (i.e. the volume corresponding tor=r_(d)=(4 Dt_(p))^(1/2)) is 0.00064 cc. This corresponds to a cube thatis 0.86 mm on a side.

This simple example demonstrates that focused ultrasound should beusable to stimulate reparative HSP's deep in the body with easilyattainable equipment:

Total ultrasound power: 5.8 watts-17 watts Pulse time 0.5 sec Pulseinterval   5 sec Total train duration (N = 10)  50 secTo expedite the treatment of larger internal volumes, a SAPRA system canbe used.

The pulsed energy source may be directed to an exterior of a body whichis adjacent to the target tissue or has a blood supply close to thesurface of the exterior of the body. Alternatively, a device may beinserted into a cavity of a body to apply the pulsed energy source tothe target tissue. Whether the energy source is applied outside of thebody or inside of the body and what type of device is utilized dependsupon the energy source selected and used to treat the target tissue.

Photostimulation, in accordance with the present invention, can beeffectively transmitted to an internal surface area or tissue of thebody utilizing an endoscope, such as a bronchoscope, proctoscope,colonoscope or the like. Each of these consist essentially of a flexibletube that itself contains one or more internal tubes. Typically, one ofthe internal tubes comprises a light pipe or multi-mode optical fiberwhich conducts light down the scope to illuminate the region of interestand enable the doctor to see what is at the illuminated end. Anotherinternal tube could consist of wires that carry an electrical current toenable the doctor to cauterize the illuminated tissue. Yet anotherinternal tube might consist of a biopsy tool that would enable thedoctor to snip off and hold on to any of the illuminated tissue.

In the present invention, one of these internal tubes is used as anelectromagnetic radiation pipe, such as a multi-mode optical fiber, totransmit the SDM or other electromagnetic radiation pulses that are fedinto the scope at the end that the doctor holds. With reference now toFIG. 18, a light generating unit 10, such as a laser having a desiredwavelength and/or frequency is used to generate electromagneticradiation, such as laser light, in a controlled, pulsed manner to bedelivered through a light tube or pipe 12 to a distal end of the scope14, illustrated in FIG. 19, which is inserted into the body and thelaser light or other radiation 16 delivered to the target tissue 18 tobe treated.

With reference now to FIG. 20, a schematic diagram is shown of a systemfor generating electromagnetic energy radiation, such as laser light,including SDM. The system, generally referred to by the reference number20, includes a laser console 22, such as for example the 810 nm nearinfrared micropulsed diode laser in the preferred embodiment. The lasergenerates a laser light beam which is passed through optics, such as anoptical lens or mask, or a plurality of optical lenses and/or masks 24as needed. The laser projector optics 24 pass the shaped light beam to adelivery device 26, such as an endoscope, for projecting the laser beamlight onto the target tissue of the patient. It will be understood thatthe box labeled 26 can represent both the laser beam projector ordelivery device as well as a viewing system/camera, such as anendoscope, or comprise two different components in use. The viewingsystem/camera 26 provides feedback to a display monitor 28, which mayalso include the necessary computerized hardware, data input andcontrols, etc. for manipulating the laser 22, the optics 24, and/or theprojection/viewing components 26.

With reference now to FIG. 21, in one embodiment, a plurality of lightbeams are generated, each of which has parameters selected so that atarget tissue temperature may be controllably raised to therapeuticallytreat the target tissue without destroying or permanently damaging thetarget tissue. This may be done, for example, by passing the laser lightbeam 30 through optics which diffract or otherwise generate a pluralityof laser light beams from the single laser light beam 30 having theselected parameters. For example, the laser light beam 30 may be passedthrough a collimator lens 32 and then through a mask 34. In aparticularly preferred embodiment, the mask 34 comprises a diffractiongrating. The mask/diffraction grating 34 produces a geometric object, ormore typically a geometric pattern of simultaneously produced multiplelaser spots or other geometric objects. This is represented by themultiple laser light beams labeled with reference number 36.Alternatively, the multiple laser spots may be generated by a pluralityof fiber optic waveguides. Either method of generating laser spotsallows for the creation of a very large number of laser spotssimultaneously over a very wide treatment field. In fact, a very highnumber of laser spots, perhaps numbering in the hundreds even thousandsor more could be simultaneously generated to cover a given area of thetarget tissue, or possibly even the entirety of the target tissue. Awide array of simultaneously applied small separated laser spotapplications may be desirable as such avoids certain disadvantages andtreatment risks known to be associated with large laser spotapplications.

Using optical features with a feature size on par with the wavelength ofthe laser employed, for example using a diffraction grating, it ispossible to take advantage of quantum mechanical effects which permitssimultaneous application of a very large number of laser spots for avery large target area. The individual spots produced by suchdiffraction gratings are all of a similar optical geometry to the inputbeam, with minimal power variation for each spot. The result is aplurality of laser spots with adequate irradiance to produce harmlessyet effective treatment application, simultaneously over a large targetarea. The present invention also contemplates the use of other geometricobjects and patterns generated by other diffractive optical elements.

The laser light passing through the mask 34 diffracts, producing aperiodic pattern a distance away from the mask 34, shown by the laserbeams labeled 36 in FIG. 21. The single laser beam 30 has thus beenformed into hundreds or even thousands of individual laser beams 36 soas to create the desired pattern of spots or other geometric objects.These laser beams 36 may be passed through additional lenses,collimators, etc. 38 and 40 in order to convey the laser beams and formthe desired pattern. Such additional lenses, collimators, etc. 38 and 40can further transform and redirect the laser beams 36 as needed.

Arbitrary patterns can be constructed by controlling the shape, spacingand pattern of the optical mask 34. The pattern and exposure spots canbe created and modified arbitrarily as desired according to applicationrequirements by experts in the field of optical engineering.Photolithographic techniques, especially those developed in the field ofsemiconductor manufacturing, can be used to create the simultaneousgeometric pattern of spots or other objects.

The present invention can use a multitude of simultaneously generatedtherapeutic light beams or spots, such as numbering in the dozens oreven hundreds, as the parameters and methodology of the presentinvention create therapeutically effective yet non-destructive andnon-permanently damaging treatment. Although hundreds or even thousandsof simultaneous laser spots could be generated and created and formedinto patterns to be simultaneously applied to the tissue, due to therequirements of not overheating the tissue, there are constraints on thenumber of treatment spots or beams which can be simultaneously used inaccordance with the present invention. Each individual laser beam orspot requires a minimum average power over a train duration to beeffective. However, at the same time, tissue cannot exceed certaintemperature rises without becoming damaged. For example, using an 810 nmwavelength laser, the number of simultaneous spots generated and usedcould number from as few as 1 and up to approximately 100 when a 0.04(4%) duty cycle and a total train duration of 0.3 seconds (300milliseconds) is used. The water absorption increases as the wavelengthis increased. For shorter wavelengths, e.g., 577 nm, the laser power canbe lower. For example, at 577 nm, the power can be lowered by a factorof 4 for the invention to be effective. Accordingly, there can be as fewas a single laser spot or up to approximately 400 laser spots when usingthe 577 nm wavelength laser light, while still not harming or damagingthe tissue.

Typically, the system of the present invention incorporates a guidancesystem to ensure complete and total retinal treatment with retinalphotostimulation. Fixation/tracking/registration systems consisting of afixation target, tracking mechanism, and linked to system operation canbe incorporated into the present invention. In a particularly preferredembodiment, the geometric pattern of simultaneous laser spots issequentially offset so as to achieve confluent and complete treatment ofthe surface.

This can be done in a controlled manner using an optical scanningmechanism 50. FIGS. 22 and 23 illustrate an optical scanning mechanism50 in the form of a MEMS mirror, having a base 52 with electronicallyactuated controllers 54 and 56 which serve to tilt and pan the mirror 58as electricity is applied and removed thereto. Applying electricity tothe controller 54 and 56 causes the mirror 58 to move, and thus thesimultaneous pattern of laser spots or other geometric objects reflectedthereon to move accordingly on the retina of the patient. This can bedone, for example, in an automated fashion using electronic softwareprogram to adjust the optical scanning mechanism 50 until completecoverage of the retina, or at least the portion of the retina desired tobe treated, is exposed to the phototherapy. The optical scanningmechanism may also be a small beam diameter scanning galvo mirrorsystem, or similar system, such as that distributed by Thorlabs. Such asystem is capable of scanning the lasers in the desired offsettingpattern.

The pattern of spots are offset at each exposure so as to create spacebetween the immediately previous exposure to allow heat dissipation andprevent the possibility of heat damage or tissue destruction. Thus, asillustrated in FIG. 24, the pattern, illustrated for exemplary purposesas a grid of sixteen spots, is offset each exposure such that the laserspots occupy a different space than previous exposures. It will beunderstood that the diagrammatic use of circles or empty dots as well asfilled dots are for diagrammatic purposes only to illustrate previousand subsequent exposures of the pattern of spots to the area, inaccordance with the present invention. The spacing of the laser spotsprevents overheating and damage to the tissue. It will be understoodthat this occurs until the entire target tissue to be treated hasreceived phototherapy, or until the desired effect is attained. This canbe done, for example, by applying electrostatic torque to amicromachined mirror, as illustrated in FIGS. 22 and 23. By combiningthe use of small laser spots separated by exposure free areas, preventsheat accumulation, and grids with a large number of spots per side, itis possible to atraumatically and invisibly treat large target areaswith short exposure durations far more rapidly than is possible withcurrent technologies.

By rapidly and sequentially repeating redirection or offsetting of theentire simultaneously applied grid array of spots or geometric objects,complete coverage of the target, can be achieved rapidly without thermaltissue injury. This offsetting can be determined algorithmically toensure the fastest treatment time and least risk of damage due tothermal tissue, depending on laser parameters and desired application.

The following has been modeled using the Fraunhoffer Approximation. Witha mask having a nine by nine square lattice, with an aperture radius 9μm, an aperture spacing of 600 μm, using a 890 nm wavelength laser, witha mask-lens separation of 75 mm, and secondary mask size of 2.5 mm by2.5 mm, the following parameters will yield a grid having nineteen spotsper side separated by 133 μm with a spot size radius of 6 μm. The numberof exposures “m” required to treat (cover confluently with small spotapplications) given desired area side-length “A”, given output patternspots per square side “n”, separation between spots “R”, spot radius “r”and desired square side length to treat area “A”, can be given by thefollowing formula:

$m = {\frac{A}{nR}{{floor}\left( \frac{R}{2r} \right)}^{2}}$

With the foregoing setup, one can calculate the number of operations mneeded to treat different field areas of exposure. For example, a 3 mm×3mm area, which is useful for treatments, would require 98 offsettingoperations, requiring a treatment time of approximately thirty seconds.Another example would be a 3 cm×3 cm area, representing the entire humanretinal surface. For such a large treatment area, a much largersecondary mask size of 25 mm by 25 mm could be used, yielding atreatment grid of 190 spots per side separated by 133 μm with a spotsize radius of 6 μm. Since the secondary mask size was increased by thesame factor as the desired treatment area, the number of offsettingoperations of approximately 98, and thus treatment time of approximatelythirty seconds, is constant.

Of course, the number and size of spots produced in a simultaneouspattern array can be easily and highly varied such that the number ofsequential offsetting operations required to complete treatment can beeasily adjusted depending on the therapeutic requirements of the givenapplication.

Furthermore, by virtue of the small apertures employed in thediffraction grating or mask, quantum mechanical behavior may be observedwhich allows for arbitrary distribution of the laser input energy. Thiswould allow for the generation of any arbitrary geometric shapes orpatterns, such as a plurality of spots in grid pattern, lines, or anyother desired pattern. Other methods of generating geometric shapes orpatterns, such as using multiple fiber optical fibers or microlenses,could also be used in the present invention. Time savings from the useof simultaneous projection of geometric shapes or patterns permits thetreatment fields of novel size, such as the 1.2 cm̂2 area to accomplishwhole-retinal treatment, in a single clinical setting or treatmentsession.

With reference now to FIG. 25, instead of a geometric pattern of smalllaser spots, the present invention contemplates use of other geometricobjects or patterns. For example, a single line 60 of laser light,formed by the continuously or by means of a series of closely spacedspots, can be created. An offsetting optical scanning mechanism can beused to sequentially scan the line over an area, illustrated by thedownward arrow in FIG. 25.

With reference now to FIG. 26, the same geometric object of a line 60can be rotated, as illustrated by the arrows, so as to create a circularfield of phototherapy. The potential negative of this approach, however,is that the central area will be repeatedly exposed, and could reachunacceptable temperatures. This could be overcome, however, byincreasing the time between exposures, or creating a gap in the linesuch that the central area is not exposed.

The field of photobiology reveals that different biologic effects may beachieved by exposing target tissues to lasers of different wavelengths.The same may also be achieved by consecutively applying multiple lasersof either different or the same wavelength in sequence with variabletime periods of separation and/or with different irradiant energies. Thepresent invention anticipates the use of multiple laser, light orradiant wavelengths (or modes) applied simultaneously or in sequence tomaximize or customize the desired treatment effects. This method alsominimizes potential detrimental effects. The optical methods and systemsillustrated and described above provide simultaneous or sequentialapplication of multiple wavelengths.

FIG. 27 illustrates diagrammatically a system which couples multipletreatment light sources into the pattern-generating optical subassemblydescribed above. Specifically, this system 20′ is similar to the system20 described in FIG. 20 above. The primary differences between thealternate system 20′ and the earlier described system 20 is theinclusion of a plurality of laser consoles, the outputs of which areeach fed into a fiber coupler 42. Each laser console may supply a laserlight beam having different parameters, such as of a differentwavelength. The fiber coupler produces a single output that is passedinto the laser projector optics 24 as described in the earlier system.The coupling of the plurality of laser consoles 22 into a single opticalfiber is achieved with a fiber coupler 42 as is known in the art. Otherknown mechanisms for combining multiple light sources are available andmay be used to replace the fiber coupler described herein.

In this system 20′ the multiple light sources 22 follow a similar pathas described in the earlier system 20, i.e., collimated, diffracted,recollimated, and directed to the projector device and/or tissue. Inthis alternate system 20′ the diffractive element must functiondifferently than described earlier depending upon the wavelength oflight passing through, which results in a slightly varying pattern. Thevariation is linear with the wavelength of the light source beingdiffracted. In general, the difference in the diffraction angles issmall enough that the different, overlapping patterns may be directedalong the same optical path through the projector device 26 to thetissue for treatment.

Since the resulting pattern will vary slightly for each wavelength, asequential offsetting to achieve complete coverage will be different foreach wavelength. This sequential offsetting can be accomplished in twomodes. In the first mode, all wavelengths of light are appliedsimultaneously without identical coverage. An offsetting steeringpattern to achieve complete coverage for one of the multiple wavelengthsis used. Thus, while the light of the selected wavelength achievescomplete coverage of the tissue, the application of the otherwavelengths achieves either incomplete or overlapping coverage of thetissue. The second mode sequentially applies each light source of avarying wavelength with the proper steering pattern to achieve completecoverage of the tissue for that particular wavelength. This modeexcludes the possibility of simultaneous treatment using multiplewavelengths, but allows the optical method to achieve identical coveragefor each wavelength. This avoids either incomplete or overlappingcoverage for any of the optical wavelengths.

These modes may also be mixed and matched. For example, two wavelengthsmay be applied simultaneously with one wavelength achieving completecoverage and the other achieving incomplete or overlapping coverage,followed by a third wavelength applied sequentially and achievingcomplete coverage.

FIG. 28 illustrates diagrammatically yet another alternate embodiment ofthe inventive system 20″. This system 20″ is configured generally thesame as the system 20 depicted in FIG. 20. The main difference residesin the inclusion of multiple pattern-generating subassembly channelstuned to a specific wavelength of the light source. Multiple laserconsoles 22 are arranged in parallel with each one leading directly intoits own laser projector optics 24. The laser projector optics of eachchannel 44 a, 44 b, 44 c comprise a collimator 32, mask or diffractiongrating 34 and recollimators 38, 40 as described in connection with FIG.21 above—the entire set of optics tuned for the specific wavelengthgenerated by the corresponding laser console 22. The output from eachset of optics 24 is then directed to a beam splitter 46 for combinationwith the other wavelengths. It is known by those skilled in the art thata beam splitter used in reverse can be used to combine multiple beams oflight into a single output. The combined channel output from the finalbeam splitter 46 c is then directed through the projector device 26.

In this system 20″ the optical elements for each channel are tuned toproduce the exact specified pattern for that channel's wavelength.Consequently, when all channels are combined and properly aligned asingle steering pattern may be used to achieve complete coverage of thetissue for all wavelengths. The system 20″ may use as many channels 44a, 44 b, 44 c, etc. and beam splitters 46 a, 46 b, 46 c, etc. as thereare wavelengths of light being used in the treatment.

Implementation of the system 20″ may take advantage of differentsymmetries to reduce the number of alignment constraints. For example,the proposed grid patterns are periodic in two dimensions and steered intwo dimensions to achieve complete coverage. As a result, if thepatterns for each channel are identical as specified, the actual patternof each channel would not need to be aligned for the same steeringpattern to achieve complete coverage for all wavelengths. Each channelwould only need to be aligned optically to achieve an efficientcombination.

In system 20″, each channel begins with a light source 22, which couldbe from an optical fiber as in other embodiments of thepattern-generating subassembly. This light source 22 is directed to theoptical assembly 24 for collimation, diffraction, recollimation anddirected into the beam splitter which combines the channel with the mainoutput.

It will be understood that the laser light generating systemsillustrated in FIGS. 20-28 are exemplary. Other devices and systems canbe utilized to generate a source of SDM laser light which can beoperably passed through to a projector device, typically in the form ofan endoscope having a light pipe or the like. Also, other forms ofelectromagnetic radiation may also be generated and used, includingultraviolet waves, microwaves, other radiofrequency waves, and laserlight at predetermined wavelengths. Moreover, ultrasound waves may alsobe generated and used to create a thermal time-course temperature spikein the target tissue sufficient to activate or produce heat shockproteins in the cells of the target tissue without damaging the targettissue itself. In order to do so, typically, a pulsed source ofultrasound or electromagnetic radiation energy is provided and appliedto the target tissue in a manner which raises the target tissuetemperature, such as between 6° C. and 11° C., transiently while only 6°C. or 1° C. or less for the long term, such as over several minutes.

It is believed that stimulating HSP production in accordance with thepresent invention can be effectively utilized in treating a wide arrayof tissue abnormalities, ailments, and even infections. For example, theviruses that cause colds primarily affect a small port of therespiratory epithelium in the nasal passages and nasopharynx. Similar tothe retina, the respiratory epithelium is a thin and clear tissue. Withreference to FIG. 29, a cross-sectional view of a human head 62 is shownwith an endoscope 14 inserted into the nasal cavity 64 and energy 16,such as laser light or the like, being directed to tissue 18 to betreated within the nasal cavity 64. The tissue 18 to be treated could bewithin the nasal cavity 64, including the nasal passages, andnasopharynx.

To assure absorption of the laser energy, or other energy source, thewavelength can be adjusted to an infrared (IR) absorption peak of water,or an adjuvant dye can be used to serve as a photosensitizer. In such acase, treatment would then consist of drinking, or topically applying,the adjuvant, waiting a few minutes for the adjuvant to permeate thesurface tissue, and then administering the laser light or other energysource 16 to the target tissue 18 for a few seconds, such as via opticalfibers in an endoscope 14, as illustrated in FIG. 29. To provide comfortof the patient, the endoscope 14 could be inserted after application ofa topical anesthetic. If necessary, the procedure could be repeatedperiodically, such as in a day or so.

The treatment would stimulate the activation or production of heat shockproteins and facilitate protein repair without damaging the cells andtissues being treated. As discussed above, certain heat shock proteinshave been found to play an important role in the immune response as wellas the well-being of the targeted cells and tissue. The source of energycould be monochromatic laser light, such as 810 nm wavelength laserlight, administered in a manner similar to that described in theabove-referenced patent applications, but administered through anendoscope or the like, as illustrated in FIG. 29. The adjuvant dye wouldbe selected so as to increase the laser light absorption. While thiscomprises a particularly preferred method and embodiment of performingthe invention, it will be appreciated that other types of energy anddelivery means could be used to achieve the same objectives inaccordance with the present invention.

With reference now to FIG. 30, a similar situation exists for the fluvirus, where the primary target is the epithelium of the upperrespiratory tree, in segments that have diameters greater than about 3.3mm, namely, the upper six generations of the upper respiratory tree. Athin layer of mucous separates the targeted epithelial cells from theairway lumen, and it is in this layer that the antigen-antibodyinteractions occur that result in inactivation of the virus.

With continuing reference to FIG. 30, the flexible light tube 12 of abronchoscope 14 is inserted through the individual's mouth 66 throughthe throat and trachea 68 and into a bronchus 70 of the respiratorytree. There the laser light or other energy source 16 is administeredand delivered to the tissue in this area of the uppermost segments totreat the tissue and area in the same manner described above withrespect to FIG. 29. It is contemplated that a wavelength of laser orother energy would be selected so as to match an IR absorption peak ofthe water resident in the mucous to heat the tissue and stimulate HSPactivation or production and facilitate protein repair, with itsattendant benefits.

With reference now to FIG. 31, a colonoscope 14 could have flexibleoptical tube 12 thereof inserted into the anus and rectum 72 and intoeither the large intestine 74 or small intestine 76 so as to deliver theselected laser light or other energy source 16 to the area and tissue tobe treated, as illustrated. This could be used to assist in treatingcolon cancer as well as other gastrointestinal issues.

Typically, the procedure could be performed similar to a colonoscopy inthat the bowel would be cleared of all stool, and the patient would lieon his/her side and the physician would insert the long, thin light tubeportion 12 of the colonoscope 14 into the rectum and move it into thearea of the colon, large intestine 74 or small intestine 76 to the areato be treated. The physician could view through a monitor the pathway ofthe inserted flexible member 12 and even view the tissue at the tip ofthe colonoscope 14 within the intestine, so as to view the area to betreated. Using one of the other fiber optic or light tubes, the tip 78of the scope would be directed to the tissue to be treated and thesource of laser light or other radiation 16 would be delivered throughone of the light tubes of the colonoscope 14 to treat the area of tissueto be treated, as described above, in order to stimulate HSP activationor production in that tissue 18.

With reference now to FIG. 32, another example in which the presentinvention can be advantageously used is what is frequently referred toas “leaky gut” syndrome, a condition of the gastrointestinal (GI) tractmarked by inflammation and other metabolic dysfunction. Since the GItract is susceptible to metabolic dysfunction similar to the retina, itis anticipated that it will respond well to the treatment of the presentinvention. This could be done by means of subthreshold, diode micropulselaser (SDM) treatment, as discussed above, or by other energy sourcesand means as discussed herein and known in the art.

With continuing reference to FIG. 32, the flexible light tube 12 of anendoscope or the like is inserted through the patient's mouth 66 throughthe throat and trachea area 68 and into the stomach 80, where the tip orend 78 thereof is directed towards the tissue 18 to be treated, and thelaser light or other energy source 16 is directed to the tissue 18. Itwill be appreciated by those skilled in the art that a colonoscope couldalso be used and inserted through the rectum 72 and into the stomach 80or any tissue between the stomach and the rectum.

If necessary, a chromophore pigment could be delivered to the GI tissueorally to enable absorption of the radiation. If, for instance,unfocused 810 nm radiation from a laser diode or LED were to be used,the pigment would have an absorption peak at or near 810 nm.Alternatively, the wavelength of the energy source could be adjusted toa slightly longer wavelength at an absorption peak of water, so that noexternally applied chromophore would be required.

It is also contemplated by the present invention that a capsuleendoscope 82, such as that illustrated in FIG. 33, could be used toadminister the radiation and energy source in accordance with thepresent invention. Such capsules are relatively small in size, such asapproximately one inch in length, so as to be swallowed by the patient.As the capsule or pill 82 is swallowed and enters into the stomach andpasses through the GI tract, when at the appropriate location, thecapsule or pill 82 could receive power and signals, such as via antenna84, so as to activate the source of energy 86, such as a laser diode andrelated circuitry, with an appropriate lens 88 focusing the generatedlaser light or radiation through a radiation-transparent cover 90 andonto the tissue to be treated. It will be understood that the locationof the capsule endoscope 82 could be determined by a variety of meanssuch as external imaging, signal tracking, or even by means of aminiature camera with lights through which the doctor would view imagesof the GI tract through which the pill or capsule 82 was passing throughat the time. The capsule or pill 82 could be supplied with its own powersource, such as by virtue of a battery, or could be powered externallyvia an antenna, such that the laser diode 86 or other energy generatingsource create the desired wavelength and pulsed energy source to treatthe tissue and area to be treated.

As in the treatment of the retina in previous applications, theradiation would be pulsed to take advantage of the micropulsetemperature spikes and associated safety, and the power could beadjusted so that the treatment would be completely harmless to thetissue. This could involve adjusting the peak power, pulse times, andrepetition rate to give spike temperature rises on the order of 10° C.,while maintaining the long term rise in temperature to be less than theFDA mandated limit of 1° C. If the pill form 82 of delivery is used, thedevice could be powered by a small rechargeable battery or over wirelessinductive excitation or the like. The heated/stressed tissue wouldstimulate activation or production of HSP and facilitate protein repair,and the attendant benefits thereof.

From the foregoing examples, the technique of the present invention islimited to the treatment of conditions at near body surfaces or atinternal surfaces easily accessible by means of fiber optics or otheroptical delivery means. The reason that the application of SDM toactivate HSP activity is limited to near surface or optically accessiblyregions of the body is that the absorption length of IR or visibleradiation in the body is very short. However, there are conditionsdeeper within tissue or the body which could benefit from the presentinvention. Thus, the present invention contemplates the use ofultrasound and/or radio frequency (RF) and even shorter wavelengthelectromagnetic (EM) radiation such as microwave which have relativelylong absorption lengths in body tissue. The use of pulsed ultrasound ispreferable to RF electromagnetic radiation to activate remedial HSPactivity in abnormal tissue that is inaccessible to surface SDM or thelike.

For deep tissue that is not near an internal orifice, a light pipe maynot be an effective means of delivering the pulsed energy. In that case,pulsed low frequency electromagnetic energy or preferably pulsedultrasound can be used to cause a series of temperature spikes in thetarget tissue.

Thus, in accordance with the present invention, a source of pulsedultrasound or electromagnetic radiation is applied to the target tissuein order to stimulate HSP production or activation and to facilitateprotein repair in the living animal tissue. In general, electromagneticradiation may be ultraviolet waves, microwaves, other radiofrequencywaves, laser light at predetermined wavelengths, etc. On the other hand,if electromagnetic energy is to be used for deep tissue targets awayfrom natural orifices, absorption lengths restrict the wavelengths tothose of microwaves or radiofrequency waves, depending on the depth ofthe target tissue. However, ultrasound is to be preferred to longwavelength electromagnetic radiation for deep tissue targets away fromnatural orifices.

The ultrasound or electromagnetic radiation is pulsed so as to create athermal time-course in the tissue that stimulates HSP production oractivation and facilitates protein repair without causing damage to thecells and tissue being treated. The area and/or volume of the treatedtissue is also controlled and minimized so that the temperature spikesare on the order of several degrees, e.g. approximately 10° C., whilemaintaining the long-term rise in temperature to be less than the FDAmandated limit, such as 1° C. It has been found that if too large of anarea or volume of tissue is treated, the increased temperature of thetissue cannot be diffused sufficiently quickly enough to meet the FDArequirements. However, limiting the area and/or volume of the treatedtissue as well as creating a pulsed source of energy accomplishes thegoals of the present invention of stimulating HSP activation orproduction by heating or otherwise stressing the cells and tissue, whileallowing the treated cells and tissues to dissipate any excess heatgenerated to within acceptable limits.

With reference now to FIG. 34, with ultrasound, a specific region deepin the body can be specifically targeted by using one or more beams thatare each focused on the target site. The pulsating heating will then belargely only in the targeted region where the beams are focused andoverlap. Pulsed ultrasound sources can also be used for abnormalities ator near surfaces as well.

As illustrated in FIG. 34, an ultrasound transducer 92 or the likegenerates a plurality of ultrasound beams 94 which are coupled to theskin via an acoustic-impedance-matching gel, and penetrate through theskin 96 and through undamaged tissue in front of the focus of the beams94 to a target organ 98, such as the illustrated liver, and specificallyto a target tissue 100 to be treated where the ultrasound beams 94 arefocused. As mentioned above, the pulsating heating will then only be atthe targeted, focused region 100 where the focused beams 94 overlap. Thetissue in front of and behind the focused region 100 will not be heatedor affected appreciably.

The present invention contemplates not only the treatment of surface ornear surface tissue, such as using the laser light or the like, deeptissue using, for example, focused ultrasound beams or the like, butalso treatment of blood diseases, such as sepsis. As indicated above,focused ultrasound treatment could be used both at surface as well asdeep body tissue, and could also be applied in this case in treatingblood. However, it is also contemplated that the SDM and similartreatment options which are typically limited to surface or near surfacetreatment of epithelial cells and the like be used in treating blooddiseases at areas where the blood is accessible through a relativelythin layer of tissue, such as the earlobe.

With reference now to FIGS. 35 and 36, treatment of blood disorderssimply requires the transmission of SDM or other electromagneticradiation or ultrasound pulses to the earlobe 102, where the SDM orother radiation source of energy could pass through the earlobe tissueand into the blood which passes through the earlobe. It would beappreciated that this approach could also take place at other areas ofthe body where the blood flow is relatively high and/or near the tissuesurface, such as fingertips, inside of the mouth or throat, etc.

With reference again to FIGS. 35 and 36, an earlobe 102 is shownadjacent to a clamp device 104 configured to transmit SDM radiation orthe like. This could be, for example, by means of one or more laserdiodes 106 which would transmit the desired frequency at the desiredpulse and pulse train to the earlobe 102. Power could be provided, forexample, by means of a lamp drive 108. Alternatively, the lamp drive 108could be the actual source of laser light, which would be transmittedthrough the appropriate optics and electronics to the earlobe 102. Theclamp device 104 would merely be used to clamp onto the patient'searlobe and cause that the radiation be constrained to the patient'searlobe 102. This may be by means of mirrors, reflectors, diffusers,etc. This could be controlled by a control computer 110, which would beoperated by a keyboard 112 or the like. The system may also include adisplay and speakers 114, if needed, for example if the procedure wereto be performed by an operator at a distance from the patient.

The proposed treatment with a train of electromagnetic or ultrasoundpulses has two major advantages over earlier treatments that incorporatea single short or sustained (long) pulse. First, the short (preferablysubsecond) individual pulses in the train activate cellular resetmechanisms like HSP activation with larger reaction rate constants thanthose operating at longer (minute or hour) time scales. Secondly, therepeated pulses in the treatment provide large thermal spikes (on theorder of 10,000) that allow the cell's repair system to more rapidlysurmount the activation energy barrier that separates a dysfunctionalcellular state from the desired functional state. The net result is a“lowered therapeutic threshold” in the sense that a lower appliedaverage power and total applied energy can be used to achieve thedesired treatment goal.

Power limitations in current micropulsed diode lasers require fairlylong exposure duration. The longer the exposure, the more important thecenter-spot heat dissipating ability toward the unexposed tissue at themargins of the laser spot. Thus, the micropulsed laser light beam of an810 nm diode laser should have an exposure envelope duration of 500milliseconds or less, and preferably approximately 300 milliseconds. Ofcourse, if micropulsed diode lasers become more powerful, the exposureduration should be lessened accordingly.

Aside from power limitations, another parameter of the present inventionis the duty cycle, or the frequency of the train of micropulses, or thelength of the thermal relaxation time between consecutive pulses. It hasbeen found that the use of a 10% duty cycle or higher adjusted todeliver micropulsed laser at similar irradiance at similar MPE levelssignificantly increase the risk of lethal cell injury. However, dutycycles of less than 10%, and preferably 5% or less demonstrate adequatethermal rise and treatment at the level of the MPE cell to stimulate abiological response, but remain below the level expected to producelethal cell injury. The lower the duty cycle, however, the exposureenvelope duration increases, and in some instances can exceed 500milliseconds.

Each micropulse lasts a fraction of a millisecond, typically between 50microseconds to 100 microseconds in duration. Thus, for the exposureenvelope duration of 300-500 milliseconds, and at a duty cycle of lessthan 5%, there is a significant amount of wasted time betweenmicropulses to allow the thermal relaxation time between consecutivepulses. Typically, a delay of between 1 and 3 milliseconds, andpreferably approximately 2 milliseconds, of thermal relaxation time isneeded between consecutive pulses. For adequate treatment, the cells aretypically exposed or hit between 50-200 times, and preferably between75-150 at each location, and with the 1-3 milliseconds of relaxation orinterval time, the total time in accordance with the embodimentsdescribed above to treat a given area which is being exposed to thelaser spots is usually less than one second, such as between 100milliseconds and 600 milliseconds on average. The thermal relaxationtime is required so as not to overheat the cells within that location orspot and so as to prevent the cells from being damaged or destroyed.While time periods of 100-600 milliseconds do not seem long, given thesmall size of the laser spots and the need to treat a relatively largearea of the target tissue, treating the entire target tissue take asignificant amount of time, particularly for a patient who is undergoingtreatment.

Other pulsed energy sources, including microwave, radio frequency andultrasound is also preferably pulsed in nature and have duty cyclesand/or pulse trains and thus lag time or intervals between micropulseenergy applications to the target tissue. Moreover, the target tissuepreviously treated with the micropulse of the energy must be allowed todissipate the heat created by the energy application in order not toexceed a predetermined upper temperature level which could permanentlydamage or even destroy the cells of the target tissue. Typically, thearea or volume of target tissue to be treated is much larger than thearea or volume of target tissue which is treated at any given moment bythe energy sources, even if multiple beams of energy are created andapplied to the target tissue.

Accordingly, the present invention may utilize the interval betweenconsecutive applications to the same location to apply energy to asecond treatment area, or additional areas, of the target tissue that isspaced apart from the first treatment area. The pulsed energy isreturned to the first treatment location, or previous treatmentlocations, within the predetermined interval of time so as to providesufficient thermal relaxation time between consecutive pulses, yet alsosufficiently treat the cells in those locations or areas properly bysufficiently increasing the temperature of those cells over time byrepeatedly applying the energy to that location in order to achieve thedesired therapeutic benefits of the invention.

It is important to return to a previously treated location within apredetermined amount of time to allow the area to cool down sufficientlyduring that time, but also to treat it within the necessary window oftime. In the case of the laser light pulsed energy applications, thelaser light is returned to the previously treated location within one tothree milliseconds, and preferably approximately two milliseconds, asone cannot wait one or two seconds and then return to a previouslytreated area that has not yet received the full treatment necessary, asthe treatment will not be as effective or perhaps not effective at all.However, during that interval of time, typically approximately 2milliseconds, at least one other area, and typically multiple areas, canbe treated with a laser light application as the laser light pulses aretypically 50 seconds to 100 microseconds in duration. This is referredto herein as microshifting. The number of additional areas which can betreated is limited only by the micopulse duration and the ability tocontrollably move the light beams from one area to another.

Currently, approximately four additional areas which are sufficientlyspaced apart from one another can be treated during the thermalrelaxation intervals beginning with a first treatment area when usinglaser light. Thus, multiple areas can be treated, at least partially,during the 200-500 millisecond exposure envelope for the first area.Thus, in a single interval of time, instead of only 100 simultaneouslight spots being applied to a treatment area, approximately 500 lightspots can be applied during that interval of time in different treatmentareas. This would be the case, for example, for a laser light beamhaving a wavelength of 810 nm. For shorter wavelengths, such as 572 nm,even a greater number of individual locations can be exposed to thelaser beams to create light spots. Thus, instead of a maximum ofapproximately 400 simultaneous spots, approximately 2,000 spots could becovered during the interval between micropulse treatments to a givenarea or location. Typically each location has between 50-200, and moretypically between 75-150, light applications applied thereto over thecourse of the exposure envelope duration (typically 200-500milliseconds) to achieve the desired treatment. In accordance with anembodiment of the present invention, the laser light would be reappliedto previously treated areas in sequence during the relaxation timeintervals for each area or location. This would occur repeatedly until apredetermined number of laser light applications to each area to betreated have been achieved.

Similarly, the one or more beams of microwave, radiofrequency and/orultrasound could be applied to second, or additional treatment areas ofthe target tissue that is spaced apart from the first treatment area,and after a predetermined interval of time returning, if necessary, tothe first treatment area of the target tissue to reapply the pulsedenergy thereto. The pulsed energy could be reapplied to a previouslytreated area in sequence during the relaxation time intervals for eacharea or location until a desired number of applications has beenachieved to each treatment area. The treatment areas must be separatedby at least a predetermined minimum distance to enable thermalrelaxation and dissipation and avoid thermal tissue damage. The pulsedenergy parameters including wavelength or frequency, duty cycle andpulse train duration are selected so as to raise the target tissuetemperature up to 11° C., such as between approximately 6°−11° C.,during application of the pulsed energy source to the target tissue toachieve a therapeutic effect, such as by stimulating HSP productionwithin the cells. However, the cells of the target tissue must be givena period of time to dissipate the heat such that the average temperaturerise of the tissue over several minutes is maintained at or below apredetermined level, such as 6° C. or less, or even 1° C. or less, overseveral minutes so as not to permanently damage the target tissue.

This is diagrammatically illustrated in FIGS. 37A-37D. FIG. 37Aillustrates with solid circles a first area having energy beams, such aslaser light beams, applied thereto as a first application. The beams arecontrollably offset or microshifted to a second exposure area, followedby a third exposure area and a fourth exposure area, as illustrated inFIG. 37B, until the locations in the first exposure area need to bere-treated by having beams applied thereto again within the thermalrelaxation time interval. The locations within the first exposure areawould then have energy beams reapplied thereto, as illustrated in FIG.37C. Secondary or subsequent exposures would occur in each exposurearea, as illustrated in FIG. 37D by the increasingly shaded dots orcircles until the desired number of exposures or hits or applications ofenergy to the target tissue area has been achieved to therapeuticallytreat these areas, diagrammatically illustrated by the blackened circlesin exposure area 1 in FIG. 37D. When a first or previous exposure areahas been completed treated, this enables the system to add an additionalexposure area, which process is repeated until the entire area to betreated has been fully treated. It should be understood that the use ofsolid circles, broken line circles, partially shaded circles, and fullyshaded circles are for explanatory purposes only, as in fact theexposure of the energy or laser light in accordance with the presentinvention is invisible and non-detectable to both the human eye as wellas known detection devices and techniques, including ophthalmoscopicallyand angiographically.

Adjacent exposure areas must be separated by at least a predeterminedminimum distance to avoid thermal tissue damage. Such distance is atleast 0.5 diameter away from the immediately preceding treated locationor area, and more preferably between 1-2 diameters away. Such spacingrelates to the actually treated locations in a previous exposure area.It is contemplated by the present invention that a relatively large areamay actually include multiple exposure areas therein which are offset ina different manner than that illustrated in FIG. 37. For example, theexposure areas could comprise the thin lines illustrated in FIGS. 25 and26, which would be repeatedly exposed in sequence until all of thenecessary areas were fully exposed and treated. In accordance with thepresent invention, the time required to treat that area to be treated issignificantly reduced, such as by a factor of 4 or 5 times, such that asingle treatment session takes much less time for the medical providerand the patient need not be in discomfort for as long of a period oftime.

In accordance with this embodiment of the invention of applying one ormore treatment beams at once, and moving the treatment beams to a seriesof new locations, then bringing the beams back to re-treat the samelocation or area repeatedly has been found to also require less powercompared to the methodology of keeping the beams in the same locationsor area during the entire exposure envelope duration. With reference toFIGS. 38-40, there is a linear relationship between the pulse length andthe power necessary, but there is a logarithmic relationship between theheat generated.

With reference to FIG. 38, a graph is provided wherein the x-axisrepresents the Log of the average power in watts of a laser and they-axis represents the treatment time, in seconds. The lower curve is forpanmacular treatment and the upper curve is for panretinal treatment.This would be for a laser light beam having a micropulse time of 50microseconds, a period of 2 milliseconds of time between pulses, andduration of train on a spot of 300 milliseconds. The areas of eachretinal spot are 100 microns, and the laser power for these 100 micronretinal spots is 0.74 watts. The panmacular area is 0.55², requiring7,000 panmacular spots total, and the panretinal area is 3.30²,requiring 42,000 laser spots for full coverage. Each RPE spot requires aminimum energy in order for its reset mechanism to be adequatelyactivated, in accordance with the present invention, namely, 38.85joules for panmacular and 233.1 joules for panretinal. As would beexpected, the shorter the treatment time, the larger the requiredaverage power. However, there is an upper limit on the allowable averagepower, which limits how short the treatment time can be.

As mentioned above, there are not only power constraints with respect tothe laser light available and used, but also the amount of power thatcan be applied to the eye without damaging eye tissue. For example,temperature rise in the lens of the eye is limited, such as between 4°C. so as not to overheat and damage the lens, such as causing cataracts.Thus, an average power of 7.52 watts could elevate the lens temperatureto approximately 4° C. This limitation in power increases the minimumtreatment time.

However, with reference to FIG. 39, the total power per pulse requiredis less in the microshift case of repeatedly and sequentially moving thelaser spots and returning to prior treated locations, so that the totalenergy delivered and the total average power during the treatment timeis the same. FIGS. 39 and 40 show how the total power depends ontreatment time. This is displayed in FIG. 39 for panmacular treatment,and in FIG. 40 for panretinal treatment. The upper, solid line or curverepresents the embodiment where there are no microshifts takingadvantage of the thermal relaxation time interval, such as described andillustrated in FIG. 24, whereas the lower dashed line represents thesituation for such microshifts, as described and illustrated in FIG. 37.FIGS. 39 and 40 show that for a given treatment time, the peak totalpower is less with microshifts than without microshifts. This means thatless power is required for a given treatment time using themicroshifting embodiment of the present invention. Alternatively, theallowable peak power can be advantageously used, reducing the overalltreatment time.

Thus, in accordance with FIGS. 38-40, a log power of 1.0 (10 watts)would require a total treatment time of 20 seconds using themicroshifting embodiment of the present invention, as described herein.It would take more than 2 minutes of time without the microshifts, andinstead leaving the micropulsed light beams in the same location or areaduring the entire treatment envelope duration. There is a minimumtreatment time according to the wattage. However, this treatment timewith microshifting is much less than without microshifting. As the laserpower required is much less with the microshifting, it is possible toincrease the power in some instances in order to reduce the treatmenttime for a given desired retinal treatment area. The product of thetreatment time and the average power is fixed for a given treatment areain order to achieve the therapeutic treatment in accordance with thepresent invention. This could be implemented, for example, by applying ahigher number of therapeutic laser light beams or spots simultaneouslyat a reduced power. Of course, since the parameters of the laser lightare selected to be therapeutically effective yet not destructive orpermanently damaging to the cells, no guidance or tracking beams arerequired, only the treatment beams as all areas can be treated inaccordance with the present invention.

Although the present invention is described for use in connection with amicropulsed laser, theoretically a continuous wave laser couldpotentially be used instead of a micropulsed laser. However, with thecontinuous wave laser, there is concern of overheating as the laser ismoved from location to location in that the laser does not stop andthere could be heat leakage and overheating between treatment areas.Thus, while it is theoretically possible to use a continuous wave laser,in practice it is not ideal and the micropulsed laser is preferred.

While the information provided in connection with graphs 38-40 isderived from observations and calculations of laser light beams as theenergy source applied to retinal eye tissue, it is believed thatapplying such pulsed light beams to other tissue will achieve similarresults in that moving the treatment beams to a series of new locations,then bringing the beams back to re-treat the same location or arearepeatedly will not only save time but also require less power comparedto the methodology of keeping the beams in the same location or areaduring the entire exposure envelope duration. Similarly, it is believedthat such power conservation will also be achieved with other sources ofpulsed energy, including coherent and non-coherent light, microwave,radiofrequency and ultrasound energy sources.

In accordance with the microshifting technique described above, theshifting or steering of the pattern of light beams may be done by use ofan optical scanning mechanism, such as that illustrated and described inconnection with FIGS. 22 and 23. For situations where the wavelength ofthe illumination or energy is much less than the distance to the volumeto be illuminated or exposed, the steering can be accomplished by usingphased arrays. The illumination or energy in this case is said to be the“far field”. Phased arrays can be used for the microwave and ultrasoundillumination application or even for the laser light beam source.

Steering for microwave, ultrasound and even for laser energy sources maybe done by use of multiple sources which provide an “array”. The basicidea for steering the illumination radiation pattern of an array isconstructive (and destructive) interference between the radiation fromthe individual members of the array of sources. With reference to FIG.41, to illustrate this, it is only necessary to consider two adjacentmembers of the array. FIG. 41 depicts the wavefront originating from twoadjacent sources.

It is evident that for a wavefront that is depicted at an angle θ withrespect to the distance a between the two sources, the amplitude of thewave from the source on the left is proportional to exp[iωt] whereas theamplitude of the wave from the source on the right is proportional toexp[iωt−kasinθ−ϕ)], where ω is the angular frequency of the radiation,and k=2π/λ.

For constructive interference, these two waves should be “in phase”,i.e.

ϕconstructive=ka sin θ+2nπ  [14]

For destructive interference, these two waves should be “out of phase”,i.e.

ϕdestructive=ka sin θ+(2n+1)π  [15]

Accordingly, the illumination will be large in the directions θ given by

sin θ=(1/ka)[ϕconstructive−2nπ]  [16]

In other words, the radiation can be steered to different desireddirections θ simply by choosing different delays ϕ.

The delays can be introduced electronically into the circuits forexciting the radiation sources. The means for doing this have also beenwell discussed in the published literature: analog delay circuits areavailable as well as digital delay circuits.

Radiation patterns for microwave, ultrasound, and laser sources arequite well-directed. If we estimate the divergence of the radiation beamfrom a source of transverse dimension 2b by the Airy disc expression

Θ½=0.6λ/b  [17]

Then at a target distance D from the source, the half-width w of theilluminated region is roughly

w=0.6λD/b  [18]

If we require the separation of the illuminated regions to be 2 w, thenthe separation of the source s is roughly 3 w:

a=1.8λD/b  [19]

This can be a small separation if the source size is chosen to be muchlarger than the radiation wavelength.

For example, for ultrasound, suppose we have a 5 MHz source with atransverse dimension of 1 cm, and suppose the desired target distance is10 cm. Then the separation distance is a≈0.5 cm.

As another example, a commercially available microwave standard gainhorn source, operating at 140-220 GHz has transverse dimensions of 13.9mm by 10.8 mm and a depth dimension of 32.2 mm. For 200 GHz, thewavelength is 0.15 cm, and for a target distance of 10 cm, the targetwidth given by the equation [18] is 1.2×0.15×10/0.6=3 cm. For thespacing a of the horns, eq. [19] then gives 9 cm.

Next, apply eqs. [17]-[19] to obtain rough estimates for a steerablearray of 810 nm laser radiation. Suppose b=2×810 nm, and suppose D=1 mm.Then eqs. [17]-[19] give η½=0.3, w=0.3 mm, and a=0.9 mm.

For the radiofrequency application, however, the wavelength of theradiofrequency radiation is typically much larger than a human bodydimensions. In that case, the treatment volume is said to be in the“near field” of the radiofrequency source. Phased arrays are not usefulin near field, and a different method of steering is required.

For radio frequency treatment, the wavelength of the radiation is muchlarger than body dimensions. Thus, for 3-6 MHz, the wavelengths rangefrom 10,000 cm to 5000 cm. Accordingly, the target region in the body isin the “near field” of the source, i.e. the target distance anddimensions are much less than the wavelength of the RF radiation. Thismeans that the relevant treatment fields are not radiation fields (asthey were in the case of microwave, ultrasound, and laser treatments),but are instead induction fields.

The induction field from an RF coil is only large over dimensionscomparable to the coil dimension. The induction magnetic fields drop offrapidly as 1/r3 for distances larger than this. Accordingly, for a coilat the surface of the body, we can picture the treatment volume asroughly a hemisphere with radius equal to that of the coil.

For coils with radii between 2 and 6 mm, the treatment volumes for thesecoils are rather close to the surface (distances comparable to the coildimensions). Larger coils can be used for deeper targets. In keepingwith the spacing criteria discussed earlier, the spacing between thecoils in a surface array would be chosen to be comparable to theindividual coil dimensions.

For the laser or other light beam and ultrasound sources, thewavelengths are much less than the distances from the sources to thetarget tissue. For these sources, then, the intensity distributions fromthe arrays can be calculated in the “far field” approximation. However,for the RF sources, the wavelength is much larger than the distancesbetween the sources to the target tissue. For these sources, theintensity distribution be calculated in the “near field” approximation.For microwaves, at high frequencies, the wavelengths are much less thanthe distance between the sources and target tissue; however, at lowmicrowave frequencies, the wavelengths can be larger than the distancebetween the sources and the target tissues. (Thus, at 1 and 100 GHz, thewavelengths are 30 cm and 3 mm, respectively). Accordingly, at highmicrowave frequencies, the “far field” approximation applies, while atlow microwave frequencies, the “near field” approximation applies.

In the far field approximation, the expressions treat kR>>1, wherek=2π/λ is the wavenumber, λ is the wavelength, and R is a typicaldistance between the source and target: In this approximation, theenergy is “radiated” from the source to the target. In the near fieldapproximation, the expressions treat kR<<1: In this approximation, thefields are not radiation fields, but are “induction” fields. The arraybehaviors are markedly different in the two approximations.

For far field “radiation” arrays, the following is taken into account.With reference now to FIG. 42, a square array of square antennas orradiation sources is shown. Each antenna has a side of length 2a, andthe shortest distance between the centers of adjacent antennas is 2d.There are a total of N antennas along a line in the x direction and Nantennas along the y-direction, for a total of N² antennas.

On using the far field approximation, we find for the intensity I_(p) ata distant observation point P:

I _(p) /I _(o)={4k ² a ⁴/(π² R _(o) ²)}Sinc²{kαa)Sinc²(kβa){Sin(Nkαd)/Sin(kαd)}²{Sin(Nkβd)/Sin(kβd)}²  [20]

In this expression, it is assumed that the observation point is locateda distance R_(o) from the antenna array and that the intensity from asingle antenna is I_(o). In addition, α and β are the deflection anglesin the x and y directions, respectively, and

Sinc(v)=Sin(v)/v  [21]

Equation [20] can also be written in terms of the coordinates X and Yalong the x and y directions in the observation plane by using theapproximate relations

α=X/R _(o)  [22]

β=Y/R _(o)  [23]

From eq. [20], we can see what the specific form of the radiationpattern from the array is. FIG. 43 is a plot of a typical radiationpattern from a square array. (Anomalies in the plot appear due to theplotting routine employed. Because of plotting inaccuracies, there israndomness in the height of some of the peaks which should not bepresent, and not all of the peaks are actually shown.) The X and Zdimensions are shown in centimeters, but these dimensions can be changedeasily in the equations below.

FIG. 44 is the form of a typical radiation pattern along the X-axis fora typical radiation pattern from a “far field” array. The patternresults from the individual features shown in FIGS. 45-47.

Specifically, it is plot of

Sinc² {k(X/R _(o))a){Sin(Nk X/R _(o))d)/Sin(k(X/R _(o))d)}²

The envelope of the pattern is determined by the Sinc²{k (X/R_(o))a)function. This is shown in FIG. 45. The width of the individual lines isdetermined by the Sin²(Nk X/R_(o))d) function. This is shown in FIG. 46.Finally, the separation of the lines is determined by the Sin²(kX/R_(o))d) in the denominator. The lines occur every time the functionhas a zero, i.e. whenever the argument of the function is some multipleof it. The function is plotted in FIG. 47.

With continuing reference to FIGS. 42-47, the widths of the individuallines and the envelope are determined by the half-widths of theSin²(Nkαd) and Sinc²(kαa) functions, respectively, and the spacingbetween the lines is determined by the zeros of the Sin²((kαd) function.

Thus, we can write directly:

Width of envelope:Δenvα=ξ(λ/a)  [24]

Spacing between lines(spots):Δ_(sep)α=λ/(2d)  [25]

Width of a single line(spot):Δ_(line)α=ξ(λ/Nd)  [26]

Number of lines(spots)alongX-axis:N=2ξ(d/a)  [27a]

Number of spots in square pattern N ²=4ξ²(d/a)²  [27b]

In these expressions ξ is a fraction on the order of ½ that describeswhere the corresponding Sinc or Sin function is about half-max. (If itis desired to observe only where these functions are larger and moreuniform in magnitude, then ξ can be chosen smaller.)

A far field array, such as that illustrated in FIG. 42, can beselectively and controllably steered. The position of the peaks can bechanged by introducing a phase delay in the excitation of the antennas.Thus, the direction in the X direction can be changed by introducing aphase delay ϕ_(n) in the nth antenna in the X-direction, that isproportional to n. To change the direction from α=0 to an arbitraryα_(o), the phase delay of the nth antenna in the X direction is

ϕ_(n)=−α_(o) nkd.  [28a]

In a similar manner the maximum peak direction in the Y direction can beshifted from β=0 to an arbitrary β_(o). To change the direction from β=0to an arbitrary β_(o), an additional phase delay is introduced to themth antenna in the Y direction is

ϕ_(m)=−β_(o) mkd  [28b]

With reference now to FIG. 48, a block diagram of its system forexciting the antennas in the array, such as that illustrated in FIG. 42,to irradiate a target tissue is shown. The array system of FIG. 48 isapplicable for the light beam, ultrasound and high frequency microwavearrays. The computer controller provides the desired power excitationand phase delays for steering the array. The computer-controlledoscillator source activates the antennas with appropriate phase delaysto steer the antenna array peaks, as described above.

A near field (induction) array, and particularly the steering of suchnear field arrays, for low frequency microwaves and RF differs markedlyfrom the far field arrays discussed above.

As an example, consider the near field (induction electric field) from acircular coil carrying an alternating current I. If the coil lies in theX-Y plane with its axis along the Z-direction, then the vector potentialA is in the azimuthal direction, and is given by

A _(ϕ)=(μl/πk)(a/ρ)^(1/2)[{1−(k ²/2)}K(k ²)−E(k ²)]  [29]

with

k ²=4aρ[(a+φ ² +Z ²]⁻¹  [30]

Here

μ is the magnetic permeability of free space

a is the radius of the current carrying coil

ρ=(X²+Y²)^(1/2)

E is the complete elliptic integral of the second kind

K is the complete elliptic integral of the first kind

The induction electric field is also in the azimuthal direction, and isgiven by

E _(ϕ) =−iωA _(ϕ)  [31]

where

ω is the angular frequency of the alternating current I.

The objective of the induction field is to heat the tissue to activateheat shock proteins. The heating is achieved by dielectric or Ohmicheating: Accordingly, the temperature rise in the tissue is proportionalto Im(∈) (ωA_(ϕ))².

FIG. 49 is a plot of (πA_(ϕ)/μl)² vs the dimensionless variables x′=X/aand z′=Z/a for a coil of radius a in the Z=0 plane with its center atX=Y=0. In the figure, x′ ranges from −3 to +3, and z′ ranges from 0 to1.

With continuing reference to FIG. 49, it is shown that the inducedtissue temperature rise drops off rapidly as the axial distance from thecoil increases. The tissue between the coil and about an axial distanceequal to the radius of the coil divided by 2 can be expected toexperience a temperature rise. Thus, if it is desired to heat a tissuethat is 5 cm from the surface, where the coil sits, the coil should beapproximately 10 cm in diameter. FIG. 49 also shows that the mainheating will occur in a circular ring equal in radius to the coilradius.

To illustrate the latter point, FIGS. 50-52 show (πA_(ϕ))/μl)² vs thedimensionless variables x′=X/a at z′=0.5 for three different coils ofthe same radius a. FIG. 50 is for a coil with its center at X=−a. FIG.51 is for a coil with its center at X=0, and FIG. 52 is for a coil withits center at X=+a. FIG. 50 illustrates (πA_(ϕ))/μl)² vs thedimensionless variable x′=X/a at z′=Z/a=0.5 for a coil of radius a inthe Z=0 plane with its center at X=−a, Y=0. FIG. 51 illustrates(πA_(ϕ))μl)² vs the dimensionless variable x′=X/a at z′=Z/a=0.5 for acoil of radius a in the Z=0 plane with its center at X=Y=0. FIG. 52illustrates (πA_(ϕ))/μl)² vs the dimensionless variable x′=X/a atz′=Z/a=0.5 for a coil of radius a in the Z=0 plane with its center atX=+a and Y=0.

FIG. 53 shows the plots of FIGS. 50-52 superimposed, where (πA_(ϕ))/μl)²vs the dimensionless variable x′=X/a at z′=Z/a=0.5 for three differentlocations of a coil of radius a in the z′=0 plane. The left-most curveis for a coil with its center at X=−a and Y=0; the middle curve is for acoil with its center at X=Y=0; and the right-most curve is for a coilwith its center at X=a and Y=0.

With continuing reference to FIGS. 49-53, with respect to depth oftreatment, if a tissue at a distance Z_(o) needs to be treated byinduction heating, a coil of radius 2Z_(o) should be used. It will treatall tissue between the surface and Z_(o). For steering, for inductionfields, the way to treat different transverse positions is not to“steer” an array by phase delay, but rather to activate individual coilssequentially. Each activated coil will treat the region below it,primarily in a circular strip beneath its circumference.

With reference now to FIG. 54, a block diagram for an induction array(near field) for RF sources and low-frequency microwave sources isshown. Here, the computer-controlled powered oscillating current sourceselects the coils sequentially in order to treat different transversetissue positions. Thus, coils 1-N are powered sequentially in order tosteer the induction fields. Thus, for the different types of radiationor energy, a different steering mechanism or system is utilized in orderto treat the desired tissue at a desired depth.

As mentioned above, the controlled manner of applying energy to thetarget tissue is intended to raise the temperature of the target tissueto therapeutically treat the target tissue without destroying orpermanently damaging the target tissue. It is believed that such heatingactivates HSPs and that the thermally activated HSPs work to reset thediseased tissue to a healthy condition, such as by removing and/orrepairing damaged proteins. It is believed by the inventors thatmaximizing such HSP activation improves the therapeutic effect on thetargeted tissue. As such, understanding the behavior and activation ofHSPs and HSP system species, their generation and activation,temperature ranges for activating HSPs and time frames of the HSPactivation or generation and deactivation can be utilized to optimizethe heat treatment of the biological target tissue.

As mentioned above, the target tissue is heated by the pulsed energy fora short period of time, such as ten seconds or less, and typically lessthan one second, such as between 100 milliseconds and 600 milliseconds.The time that the energy is actually applied to the target tissue istypically much less than this in order to provide intervals of time forheat relaxation so that the target tissue does not overheat and becomedamaged or destroyed. For example, as mentioned above, laser lightpulses may last on the order of microseconds with several millisecondsof intervals of relaxed time.

Thus, understanding the sub-second behaviors of HSPs can be important tothe present invention. The thermal activation of the HSPs in SDM istypically described by an associated Arrhenius integral,

Ω=∫dt Aexp[−E/k _(B) T(t)]  [28]

where the integral is over the treatment time and

A is the Arrhenius rate constant for HSP activation

E is the activation energy

T(t) is the temperature of the thin RPE layer, including thelaser-induced temperature rise

The laser-induced temperature rise—and therefore the activationArrhenius integral—depends on both the treatment parameters (e.g., laserpower, duty cycle, total train duration) and on the RPE properties(e.g., absorption coefficients, density of HSPs). It has been foundclinically that effective SDM treatment is obtained when the Arrheniusintegrals is of the order of unity.

The Arrhenius integral formalism only takes into account a forwardreaction, i.e. only the HSP activation reaction): It does not take intoaccount any reverse reactions in which activated HSPs are returned totheir inactivated states. For the typical subsecond durations of SDMtreatments, this appears to be quite adequate. However, for longerperiods of time (e.g. a minute or longer), this formalism is not a goodapproximation: At these longer times, a whole series of reactions occursresulting in much smaller effective HSP activation rates. This is thecase during the proposed minute or so intervals between SDM applicationsin the present invention disclosure.

In the published literature, the production and destruction of heatshock proteins (HSPs) in cells over longer durations is usuallydescribed by a collection of 9-13 simultaneous mass-balance differentialequations that describe the behavior of the various molecular speciesinvolved in the life cycle of an HSP molecule. These simultaneousequations are usually solved by computer to show the behavior in time ofthe HSPs and the other species after the temperature has been suddenlyraised.

These equations are all conservation equations based on the reactions ofthe various molecular species involved in the activity of HSPs. Todescribe the behavior of the HSPs in the minute or so intervals betweenrepeated applications of SDM, we shall use the equations described in M.Rybinski, Z. Szymanska, S. Lasota, A. Gambin (2013) Modeling theefficacy of hyperthermia treatment. Journal of the Royal SocietyInterface 10, No. 88, 20130527 (Rybinski et al (2013)). The speciesconsidered in Rybinski et al (2013) are shown in Table 4.

TABLE 4 HSP system species in Rybinski et al (2013) description: HSPubiquitous heat shock protein of molecular weight 70 Da (in free,activated state) HSF heat shock (transcription) factor that has no DNAbinding capability HSF₃ (trimer) heat shock factor capable of binding toDNA, formed from HSF HSE heat shock element, a DNA site that initiatestranscription of HSP when bound to HSF₃ mRNA messenger RNA molecule forproducing HSP S substrate for HSP binding: a damaged protein P properlyfolded protein HSP.HSF a complex of HSP bound to HSF (unactivated HSPs)HSF₃.HSE a complex of HSF₃ bound to HSE, that induces transcription andthe creation of a new HSP mRNA molecule HSP.S a complex of HSP attachedto damaged protein (HSP actively repairing the protein)

The coupled simultaneous mass conservation equations for these 10species are summarized below as eqs. [29]-[38]:

d[HSP]/dt=(l ₁ +k ₁₀)[HSPS]+l ₂ [HSPHSF]+k ₄ [mRNA]−k ₁ [S][HSP]−k ₂[HSP][HSF]−l ₃ [HSP][HSF ₃ ]−k ₉ [HSP]  [29]

d{HSF]/dt=l ₂ [HSPHSF]+2l ₃ [HSP][HSF ₃ ]+k ₆ [HSPHSF][S]−k ₂[HSP][HSF]−3k ₃ [HSF] ³ −l ₆ [HSPS][HSF]  [30]

d[S]/dt=k ₁₁ {[P]+l ₁ [HSPS]+l ₆ [SPS][HSF]−k ₁ [S][HSP]−k ₆[HSPHSF][S]  [31]

d[HSPHSF]/dt=k ₂ [HSP][HSF]+l ₆ [HSPS][HSF]+l ₃ [HSP][HSF ₃ ]l ₂[HSPHSF]−k ₆ [HSPHSF][S]  [32]

d[HSPS]/dt=k ₁ [S][HSP]+k ₆ [HSPHSF][5]−(l ₁ +k ₁₀)[HSPS]−l ₆[HSPS][HSF]  [33]

d[HSF ₃ ]/dt=k ₃ [HSF] ³ +l ₇ [HSF ₃ ][HSE]−l ₃ [HSP][HSF ₃ ]−k ₇ [HSF ₃][HSE]  [34]

d[HSE]/dt=l ₇ [HSF ₃ ][HSE]−k ₇ [HSF ₃ ][HSE]  [35]

d[HSF ₃ HSE]/dt=k ₇ [HSF ₃ ][HSE]−l ₇ [HSF ₃ ][HSE]  [36]

d[mRNA]/dt=k ₈ [HSF ₃ HSE]−k ₅ [mRNA]  [37]

d[P]/dt=k ₁₀ [HSPS]−k ₁₁ [P]  [38]

In these expressions, [ ] denotes the cellular concentration of thequantity inside the bracket. For Rybinski et al (2013), the initialconcentrations at the equilibrium temperature of 310K are given in Table5.

TABLE 5 Initial values of species at 310K for a typical cell inarbitrary units [Rybinski et al (2013)]. The arbitrary units are chosenby Rybinski et al for computational convenience: to make the quantitiesof interest in the range of 0.01-10. [HSP(0)] 0.308649 [HSF(0)] 0.150836[S(0)] 0.113457 [HSPHSF(0)] 2.58799 [HSPS(0)] 1.12631 [HSF₃(0)]0.0444747 [HSE(0)] 0.957419 [HSF₃HSE(0)] 0.0425809 [mRNA(0)] 0.114641[P(0)] 8.76023

The Rybinski et al (2013) rate constants are shown in Table 6.

TABLE 6 Rybinski et al (2013) rate constants giving rates in min⁻¹ forthe arbitrary concentration units of the previous table. l₁ = 0.0175 k₁= 1.47 l₂ = 0.0175 k₂ = 1.47 l₃ = 0.020125 k₃ = 0.0805 k₄ = 0.1225 k₅ =0.0455 k₆ = 0.0805 l₆ = 0.00126 k₇ = 0.1225 l₇ = 0.1225 k₈ = 0.1225 k₉ =0.0455 k₁₀ = 0.049 k₁₁ = 0.00563271

The initial concentration values of Table 5 and the rate constants ofTable 6 were determined by Rybinski et al (2013) to correspond toexperimental data on overall HSP system behavior when the temperaturewas increased on the order of 5° C. for several (e.g. 350) minutes.

Note that the initial concentration of HSPs is100×0.308649/(8.76023+0.113457+1.12631)}=3.09% of the total number ofproteins present in the cell.

Although the rate constants of Table 6 are used by Rybinski et al forT=310+5+31 5K, it is likely that very similar rate constants exist atother temperatures. In this connection, the qualitative behavior of thesimulations is similar for a large range of parameters. For convenience,we shall assume that the values of the rate constants in Table 6 are agood approximation for the values at the equilibrium temperature ofT=310K.

The behavior of the different components in the Rybinski et al cell isdisplayed in FIGS. 55A and 55B for 350 minutes for the situation wherethe temperature is suddenly increased 5K at t=0 from an ambient 310K.

With continuing reference to FIG. 55, the behavior of HSP cellularsystem components during 350 minutes following a sudden increase intemperature from 37° C. to 42° C. is shown.

Here, the concentrations of the components are presented incomputationally convenient arbitrary units. S denotes denatured ordamaged proteins that are as yet unaffected by HSPs; HSP denotes free(activated) heat shock proteins; HSP:S denotes activated HSPs that areattached to the damaged proteins and performing repair; HSP:HSF denotes(inactive) HSPs that are attached to heat shock factor monomers; HSFdenotes a monomer of heat shock factor; HSF₃ denotes a trimer of heatshock factor that can penetrate the nuclear membrane to interact with aheat shock element on the DNA molecule; HSE:HSF₃ denotes a trimer ofheat shock factor attached to a heat shock element on the DNA moleculethat initiates transcription of a new mRNA molecule; mRNA denotes themessenger RNA molecule that results from the HSE:HSF₃, and that leads tothe production of a new (activated) HSP molecule in the cell'scytoplasm.

FIG. 55 shows that initially the concentration of activated HSPs is theresult of release of HSPs sequestered in the molecules HSPHSF in thecytoplasm, with the creation of new HSPs from the cell nucleus via mRNAnot occurring until 60 minutes after the temperature rise occurs. FIG.55 also shows that the activated HSPs are very rapidly attached todamaged proteins to begin their repair work. For the cell depicted, thesudden rise in temperature also results in a temporary rise in damagedprotein concentration, with the peak in the damaged proteinconcentration occurring about 30 minutes after the temperature increase.

FIG. 55 shows what the Rybinski et al equations predict for thevariation of the 10 different species over a period of 350 minutes.However, the present invention is concerned with SDM application is onthe variation of the species over the much shorter O(minute) intervalbetween two applications of SDM at any single retinal locus. It will beunderstood that the preferred embodiment of SDM in the form of laserlight treatment is analyzed and described, but it is applicable to othersources of energy as well.

With reference now to FIGS. 56A-56H, the behavior of HSP cellular systemcomponents during the first minute following a sudden increase intemperature from 37° C. to 42° C. using the Rybinski et al. (2013)equations with the initial values and rate constants of Tables 5 and 6are shown. The abscissa denotes time in minutes, and the ordinate showsconcentration in the same arbitrary units as in FIG. 56.

FIG. 56 shows that the nuclear source of HSPs plays virtually no roleduring a 1 minute period, and that the main source of new HSPs in thecytoplasm arises from the release of sequestered HSPs from the reservoirof HSPHSF molecules. It also shows that a good fraction of the newlyactivated HSPs attach themselves to damaged proteins to begin the repairprocess.

The initial concentrations in Table 5 are not the equilibrium values ofthe species, i.e. they do not give d[ . . . ]/dt=0, as evidenced by thecurves in FIGS. 55 and 56. The equilibrium values that give d[ . . .]/dt=0 corresponding to the rate constants of Table 6 are found to bethose listed in Table 7.

TABLE 7 Equilibrium values of species in arbitrary units [Rybinski et al(2013)] corresponding to the rate constants of Table 6. The arbitraryunits are those chosen by Rybinski et al for computational convenience:to make the quantities of interest in the range of 0.01-10. [HSP(equil)]0.315343 [HSF(equil)] 0.255145 [S(equil)] 0.542375 [HSPHSF(equil)]1.982248 [HSPS(equil)] 5.05777 [HSF₃(equil)] 0.210688 [HSE(equil)]0.206488 [HSF₃HSE(equil)] 0.643504 [mRNA(equil)] 0.1171274 [P(equil)]4.39986

Note that the equilibrium concentration of HSPs is100×{0.315343/(4.39986+5.05777+0.542375)}=3.15% of the total number ofproteins present in the cell. This is comparable, but less than theanticipated 5%-10% total number of proteins found by other researchers.However, we have not attempted to adjust percentage upwards expectingthat the general behavior will not be appreciably changed as indicatedby other researchers.

The inventors have found that a first treatment to the target tissue maybe performed by repeatedly applying the pulsed energy (e.g., SDM) to thetarget tissue over a period of time so as to controllably raise atemperature of the target tissue to therapeutically treat the targettissue without destroying or permanently damaging the target tissue. A“treatment” comprises the total number of applications of the pulsedenergy to the target tissue over a given period of time, such as dozensor even hundreds of light or other energy applications to the targettissue over a short period of time, such as a period of less than tenseconds, and more typically a period of less than one second, such as100 milliseconds to 600 milliseconds. This “treatment” controllablyraises the temperature of the target tissue to activate the heat shockproteins and related components.

What has been found, however, is that if the application of the pulsedenergy to the target tissue is halted for an interval of time, such asan interval of time that exceeds the first period of time comprising the“first treatment”, which may comprise several seconds to severalminutes, such as three seconds to three minutes or more preferably tenseconds to ninety seconds, and then a second treatment is performed onthe target tissue after the interval of time within a single treatmentsession or office visit, wherein the second treatment also entailsrepeatedly reapplying the pulsed energy to the target tissue so as tocontrollably raise the temperature of the target tissue totherapeutically treat the target tissue without destroying orpermanently damaging the target tissue, the amount of activated HSPs andrelated components in the cells of the target tissue is increasedresulting in a more effective overall treatment of the biologicaltissue. In other words, the first treatment creates a level of heatshock protein activation of the target tissue, and the second treatmentincreases the level of heat shock protein activation in the targettissue above the level due to the first treatment. Thus, performingmultiple treatments to the target tissue of the patient within a singletreatment session or office visit enhances the overall treatment of thebiological tissue so long as the second or additional treatments areperformed after an interval of time which does not exceed several minutebut which is of sufficient length so as to allow temperature relaxationso as not to damage or destroy the target tissue.

This technique may be referred to herein as “stair-stepping” in that thelevels of activated HSP production increase with the subsequenttreatment or treatments within the same office visit treatment session.This “stair-stepping” technique may be described by a combination of theArrhenius integral approach for subsecond phenomena with the Rybinski etal. (2013) treatment of intervals between repeated subsecondapplications of the SDM or other pulsed energy.

For the proposed stair-stepping SDM (repetitive SDM applications)proposed in this invention disclosure, there are some importantdifferences from the situation depicted in FIG. 55:

-   -   SDM can be applied prophylactically to a healthy cell, but        oftentimes SDM will be applied to a diseased cell. In that case,        the initial concentration of damaged proteins [S(0)] can be        larger than given in Table 7. We shall not attempt to account        for this, assuming that the qualitative behavior will not be        changed.    -   The duration of a single SDM application is only subseconds,        rather than the minutes shown in FIG. 55. The Rybinski et al        rate constants are much smaller than the Arrhenius constants:        the latter give Arrhenius integrals of the order of unity for        subsecond durations, whereas the Rybinski et al rate constants        are too small to do that. This is an example of the different        effective rate constants that exist when the time scales of        interest are different: The Rybinski et al rate constants apply        to phenomena occurring over minutes, whereas the Arrhenius rate        constants apply to subsecond phenomena.

Accordingly, to analyze what happens in the proposed stair-stepping SDMtechnique for improving the efficacy of SDM, we shall combine theArrhenius integral treatment appropriate for the subsecond phenomenawith the Rybinski et al (2013) treatment appropriate for the phenomenaoccurring over the order of a minute interval between repeated SDMapplications:

-   -   SDM subsecond application described by Arrhenius integral        formalism    -   Interval of O(minute) between SDM applications described by        Rybinski et al (2013) equations

Specifically, we consider two successive applications of SDM, each SDMmicropulse train having a subsecond duration.

-   -   For the short subsecond time scale, we assume that the        unactivated HSP's that are the source of the activated (free)        HSP's are all contained in the HSPHSF molecules in the        cytoplasm. Accordingly, the first SDM application is taken to        reduce the cytoplasmic reservoir of unactivated HSPs in the        initial HSPHSF molecule population from

[HSPHSF(equil)] to [HSPHSF(equil)]exp[−Ω],

-   -   and to increase the initial HSP molecular population from

[HSP(equil)] to [HSP(equil)]+[HSPHSF(equil)](1−exp[−Ω])

-   -   as well as to increase the initial HSF molecular population from

[HSF(equil)] to [HSF(equil)]+[HSPHSF(equil)](1−exp[−Ω])

-   -   The equilibrium concentrations of all of the other species will        be assumed to remain the same after the first SDM application    -   The Rybinski et al equations are then used to calculate what        happens to [HSP] and [HSPHSF] in the interval Xt=O(minute)        between the first SDM application and the second SDM        application, with the initial values of HSP, HSF and HSPHSF        after the first SDM application taken to be

[HSP(SDM1)]=[HSP(equil)]+[HSPHSF(equil)](1−exp[−Ω])

[HSF(SDM1)]=[HSF(equil)]+[HSPHSF(equil)](1−exp[−Ω])

and

[HSPHSF(SDM1)]=[HSPHSF(equil)]exp[−Ω]

-   -   For the second application of SDM after the interval Xt, the        values of [HSP], [HSF] and {HSPHSF] after the SDM will be taken        to be

[HSP(SDM2)]=[HSP(λt)]+[HSPHSF(λt)](1−exp[−Ω])

[HSF(SDM2)]=[HSF(λt)]+[HSPHSF(λt)](1−exp[−Ω])

and

[HSPHSF(SDM2)]=[HSPHSF(λt)]exp[−Ω]

-   -   where [HSP(λt)], [HSF(λt)], and [HSPHSF(λt)] are the values        determined from the Rybinski et al (2013) equations at the time        Xt.    -   Our present interest is in comparing [HSP[SDM2)] with        [HSP[SDM1)], to see if the repeated application of SDM at an        interval λt following the first application of SDM has resulted        in more activated (free) HSP's in the cytoplasm. The ratio β(λt,        Ω)=[HSP(SDM2)]/[HSP(SDM1)]={[{[HSP(λt)]+[HSPHSF(λt)](1−exp[−0])}/{[HSP(0)]+[HSPHSF(0)](1−exp[−0])}    -   provides a direct measure of the improvement in the degree of        HSP activation for a repeated application of SDM after an        interval λt from the first SDM application.

The HSP and HSPHSF concentrations can vary quite a bit in the intervalλt between SDM applications.

FIGS. 57A and 57B illustrate the variation in the activatedconcentrations [HSP] and the unactivated HSP in the cytoplasmicreservoir [HSPHSF] during an interval λt=1 minute between SDMapplications when the SDM Arrhenius integral 0=1 and the equilibriumconcentrations are as given in Table 7.

Although only a single repetition (one-step) is treated here, it isapparent that the procedure could be repeated to provide a multiplestair-stepping events as a means of improving the efficacy of SDM, orother therapeutic method involving activation of tissue HSPs.

Effects of varying the magnitude of the Arrhenius integral Ω andinterval λt between two distinct treatments separated by an interval oftime are shown by the following examples and results.

Nine examples generated with the procedure described above are presentedin the following. All of the examples are of a treatment consisting oftwo SDM treatments, with the second occurring at a time λt following thefirst, and they explore:

-   -   The effect of different magnitude Arrhenius integrals Ω in the        SDM treatments [Three different Ω's are considered: 0=0.2, 0.5        and 1.0]    -   The impact of varying the interval λt between the two SDM        treatments [Three different λt's are considered: λt=15 sec., 30        sec., and 60 sec.

As indicated above, the activation Arrhenius integral Ω depends on boththe treatment parameters (e.g., laser power, duty cycle, total trainduration) and on the RPE properties (e.g., absorption coefficients,density of HSPs).

Table 8 below shows the effect of different Ω(Ω=0.2, 0.5, 1) on the HSPcontent of a cell when the interval between the two SDM treatments isλt=1 minute. Here the cell is taken to have the Rybinski et al (2013)equilibrium concentrations for the ten species involved, given in Table7.

Table 8 shows four HSP concentrations (in the Rybinski et al arbitraryunits) each corresponding to four different times:

-   -   Before the first SDM treatment: [HSP(equil)]    -   Immediately after the first SDM application: [HSP(SDM1)]    -   At the end of the interval λt following the first SDM treatment:        [HSP(λt)]    -   Immediately after the second SDM treatment at λt: [HSP(SDM2)]    -   Also shown is the improvement factor over a single treatment:        β=[HSP(SDM2)]/[HSP(SDM1)]

TABLE 8 HSP concentrations at the four times just described in the text:Effect of varying the SDM Ω for two SDM applications on a cell when thetreatments are separated by  

 t = 0.25 minutes = 15 seconds. [HSP_((equil))] [HSP_((SDM1))] [HSP( 

 t)] [HSP_((SDM2))] β Ω = 0.2 0.315 0.67 0.54 0.95 1.27 Ω = 0.5 0.3151.10 0.77 1.34 1.22 Ω = 1.0 0.315 1.57 0.93 1.71 1.09

Table 9 is the same as Table 8, except that it is for an intervalbetween SDM treatments of λt=0.5 minutes=30 seconds.

TABLE 9 HSP concentrations at the four times described in the text:Effect of varying the SDM Ω for two SDM treatments on a cell when thetreatments are separated by 

 t = 0.5 minutes = 30 seconds. [HSP_((equil))] [HSP_((SDM1))] [HSP( 

 t)] [HSP_((SDM2))] β Ω = 0.2 0.315 0.67 0.44 0.77 1.14 Ω = 0.5 0.3151.10 0.58 1.18 1.08 Ω = 1.0 0.315 1.57 0.67 1.59 1.01

Table 10 is the same as the Tables 8 and 9, except that the treatmentsare separated by one minute, or sixty seconds.

TABLE 10 HSP concentrations at the four times just described in thetext: Effect of varying the SDM Ω for two SDM treatments on a normal(healthy) cell when the treatments are separated by 

 t = 1 minute = 60 seconds. [HSP_((equil))] [HSP_((SDM1))] [HSP( 

 t)] [HSP_((SDM2))] β Ω = 0.2 0.315 0.67 0.30 0.64 0.95 Ω = 0.5 0.3151.10 0.37 1.06 0.96 Ω = 1.0 0.315 1.57 0.48 1.51 0.96

Tables 8-10 show that:

-   -   The first treatment of SDM increases [HSP] by a large factor for        all three Ω's, although the increase is larger the larger Ω.        Although not displayed explicitly in the tables, the increase in        [HSP] comes at the expense of the cytoplasmic reservoir of        sequestered (unactivated) HSP's: [HSPHSF(SDM1)] is much smaller        than [HSPHSF(equil)]    -   [HSP] decreases appreciably in the interval λt between the two        SDM treatments, with the decrease being larger the larger λt is.        (The decrease in [HSP] is accompanied by an increase in both        [HSPHSF]— as shown in FIG. 44 and in [HSPS] during the interval        λt—indicating a rapid replenishment of the cytoplasmic reservoir        of unactivated HSP's and a rapid attachment of HSP's to the        damaged proteins.)    -   For λt less than 60 seconds, there is an improvement in the        number of activated (free) HSP's in the cytoplasm for two SDM        treatments rather than a single treatment.    -   The improvement increases as λt becomes smaller.    -   For λt becoming as large as 60 seconds, however, the ratio        β=[HSP(SDM2)]/[HSP(SDM1)] becomes less than unity, indicating no        improvement in two SDM treatments compared to a single SDM        treatment although this result can vary depending on energy        source parameters and tissue type that is treated.    -   The improvement for λt<60 seconds is larger the smaller the SDM        Arrhenius integral Ω is.

The results for the improvement ratio β=[HSP(SDM2)]/[HSP(SDM1)] aresummarized in FIG. 45, where the improvement ratioβ=[HSP(SDM2)]/[HSP(SDM1)]vs. interval between SDM treatments λt (inseconds) for three values of the SDM Arrhenius integral Ω, and for thethree values of the interval λt=15 sec, 30 sec, and 60 sec. Theuppermost curve is for Ω=0.2; the middle curve is for Ω=0.5; and thebottom curve is for Ω=1.0. These results are for the Rybinski et al(2013) rate constants of Table 6 and the equilibrium speciesconcentrations of Table 4.

It should be appreciated that results of Tables 8-10 and FIG. 58 are forthe Rybinski et al. (2013) rate constants of Table 6 and the equilibriumconcentrations of Table 7. The actual concentrations and rate constantsin a cell may differ from these values, and thus the number results inTables 8-10 and FIG. 58 should be taken as representative rather thanabsolute. However, they are not anticipated to be significantlydifferent. Thus, performing multiple intra-sessional treatments on asingle target tissue location or area, such as a single retinal locus,with the second and subsequent treatments following the first after aninterval anywhere from three seconds to three minutes, and preferablyten seconds to ninety seconds, should increase the activation of HSPsand related components and thus the efficacy of the overall treatment ofthe target tissue. The resulting “stair-stepping” effect achievesincremental increases in the number of heat shock proteins that areactivated, enhancing the therapeutic effect of the treatment. However,if the interval of time between the first and subsequent treatments istoo great, then the “stair-stepping” effect is lessened or not achieved.

The technique of the present invention is especially useful when thetreatment parameters or tissue characteristics are such that theassociated Arrhenius integral for activation is low, and when theinterval between repeated applications is small, such as less thanninety seconds, and preferably less than a minute. Accordingly, suchmultiple treatments must be performed within the same treatment session,such as in a single office visit, where distinct treatments can have awindow of interval of time between them so as to achieve the benefits ofthe technique of the present invention.

Although several embodiments have been described in detail for purposesof illustration, various modifications may be made without departingfrom the scope and spirit of the invention. Accordingly, the inventionis not to be limited, except as by the appended claims.

What is claimed is:
 1. A process for heat treating biological tissue,comprising the steps of: providing a plurality of energy emitters formedinto an array; generating treatment energy from the plurality ofemitters; and applying the treatment energy to target tissue; whereinthe treatment energy has energy and application parameters selected soas to raise the target tissue temperature sufficiently to create atherapeutic effect while maintaining an average temperature of thetarget tissue over several minutes at or below a predeterminedtemperature so as to not destroy or permanently damage the targettissue.
 2. The process of claim 1, wherein the selected energy andapplication parameters comprise tissue application spot size or area,average power or average power density, and exposure duration.
 3. Theprocess of claim 1, including the steps of: providing an initialtreatment to the target tissue by repeatedly applying the energy to thetarget tissue in a pulsed manner for the exposure duration comprisingless than one second; halting application of the pulsed energy to thetarget tissue for an interval of time comprising between three secondsand three minutes; and providing a secondary treatment to the targettissue after the interval of time, within a single treatment session, byrepeatedly reapplying the pulsed energy for the exposure durationcomprising less than one second.
 4. The process of claim 1, whereinduring an interval of time, comprising less than one second, betweenapplications of energy applied to a first area of the target tissue,applying treatment energy to a second area of the target tissuesufficiently spaced apart from the first area of the target tissue toavoid thermal tissue damage of the target tissue.
 5. The process ofclaim 4, wherein repeatedly applying, in an alternating manner duringthe same treatment session, the treatment energy to each of the firstand second areas of the target tissue until a predetermined number ofenergy applications to each of the first and second areas of the targettissue has been achieved.
 6. The process of claim 4, including the stepof introducing a phase delay in the activation of the energy emitters ofthe array to generate treatment energy in a phased manner using apredetermined delay of activation in order to apply treatment energy toeach of the first and second areas of the target tissue.
 7. The processof claim 4, including the step of activating the energy emitters of thearray sequentially in order to apply treatment energy to each of thefirst and second areas of the target tissue.
 8. The process of any ofclaims 1-4, wherein the treatment energy raises the target tissue to upto eleven degrees Celsius at least during application of the pulsedtreatment energy thereto.
 9. The process of any of claims 1-4, whereinthe average target tissue temperature is maintained at six degreesCelsius or less over several minutes.
 10. The process of claim 9,wherein the average target tissue temperature is maintained at onedegree Celsius or less over several minutes.
 11. The process of any ofclaims 1-4, wherein the applying step comprises the step of stimulatingheat shock protein activation in the target tissue.
 12. The process ofany of claims 1-4, wherein the treatment energy and applicationparameters are selected to have an average power density of 100-590watts/square centimeter of target tissue, a spot size between 100-500microns, and a train exposure duration of 500 milliseconds or less. 13.The process of any of claims 1-4, wherein the treatment energy compriseslight energy, radio frequency energy, microwave energy or ultrasoundenergy.
 14. The process of claim 13, wherein the treatment energycomprises light beams having a wavelength between 570 nm and 1300 nm anda duty cycle of less than 10%.
 15. The process of claim 13, wherein thetreatment energy comprises a radio frequency between approximately threeto six megahertz, a duty cycle of between approximately 2.5% and 5%. 16.The process of claim 15, wherein the radio frequency is generated withcoils having a radii between approximately 2 mm and 10 cm and betweenapproximately 13 and 57 amp turns.
 17. The process of claim 13, whereinthe treatment energy comprises a microwave frequency betweenapproximately 10 to 20 GHz and a duty cycle between approximately 2% to5% and an average power of approximately 8 and 52 watts.
 18. The processof claim 13, wherein the treatment energy comprises ultrasound having afrequency between approximately 1 MHz and 5 MHz, a duty cycle betweenapproximately 2% to 10% and a power of between approximately 0.46 and28.6 watts.